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Development of methoxy poly(ethylene glycol)-block-poly(caprolactone) amphiphilic diblock copolymer nanoparticulate… Letchford, Kevin John 2008

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DEVELOPMENT OF METHOXY POLY(ETHYLENE GLYCOL)- BLOCK-POLY(CAPROLACTONE) AMPHIPHILIC DIBLOCK COPOLYMER NANOPARTICULATE FORMULATIONS FOR THE DELIVERY OF PACLITAXEL  by  Kevin John Letchford B.Sc. (Chemistry), University of British Columbia, 1997 B.Sc. (Pharmaceutical Sciences), University of British Columbia, 2000  A THESIS SUBMITTED IN PARTIAL FULFILLMENT OF THE REQUIREMENT FOR THE DEGREE OF  DOCTOR OF PHILOSOPHY  in  The Faculty of Graduate Studies (Pharmaceutical Sciences) THE UNIVERSITY OF BRITISH COLUMBIA (Vancouver)    October 2008        ! Kevin John Letchford 2008  ii ABSTRACT The goal of this project was to develop a non-toxic amphiphilic diblock copolymer nanoparticulate drug delivery system that will solubilize paclitaxel (PTX) and retain the drug in plasma. Methoxy poly(ethylene glycol)-block-poly("-caprolactone) (MePEG-b- PCL) diblock copolymers loaded with PTX were characterized and their physicochemical properties were correlated with their performance as nanoparticulate drug delivery systems. A series of MePEG-b-PCL was synthesized with PCL blocks ranging from 2- 104 repeat units and MePEG blocks of 17, 44 or 114 repeat units. All copolymers were water soluble and formed micelles except MePEG114-b-PCL104, which was water insoluble and formed nanospheres.  Investigation of the effects of block length on the physicochemical properties of the nanoparticles was used to select appropriate copolymers for development as PTX nanoparticles. The critical micelle concentration, pyrene partition coefficient and diameter of nanoparticles were found to be dependent on the PCL block length. Copolymers based on a MePEG molecular weight of 750 g/mol were found to have temperature dependent phase behavior.  Relationships between the concentration of micellized drug and the compatibility between the drug and core-forming block, as determined by the Flory-Huggins interaction parameter, and PCL block length were developed. Increases in the compatibility between PCL and the drug, as well as longer PCL block lengths resulted in increased drug solubilization.   iii The physicochemical properties and drug delivery performance characteristics of MePEG114-b-PCL19 micelles and MePEG114-b-PCL104 nanospheres were compared. Nanospheres were larger, had a more viscous core, solubilized more PTX and released it slower, compared to micelles. No difference was seen in the hemocompatibility of the nanoparticles as assessed by plasma coagulation time and erythrocyte hemolysis. Micellar PTX had an in vitro plasma distribution similar to free drug. The majority of micellar PTX associated with the lipoprotein deficient plasma fraction (LPDP). In contrast, nanospheres were capable of retaining more of the encapsulated drug with significantly less PTX partitioning into the LPDP fraction.  In conclusion, although both micelles and nanospheres were capable of solubilizing PTX and were hemocompatible, PTX nanospheres may offer the advantage of prolonged blood circulation, based on the in vitro plasma distribution data, which showed that nanospheres retained PTX more effectively.  iv TABLE OF CONTENTS  Abstract.............................................................................................................................. ii Table of Contents ............................................................................................................. iv List of Tables ................................................................................................................... vii List of Figures.................................................................................................................viii List of Abbreviations ....................................................................................................... xi Acknowledgements ......................................................................................................... xv Co-Authorship Statement ............................................................................................ xvii  Chapter 1 Introduction ................................................................................................... 1 1.1 Project Overview ................................................................................................... 1 1.2 Nanoparticulate Drug Delivery Systems ............................................................... 4 1.2.1 Definitions and Classification......................................................................... 4 1.2.2 Therapeutic Applications ................................................................................ 8 1.2.3 Amphiphilic Block Copolymer Drug Delivery Systems .............................. 12 1.2.3.1 Amphiphilic Block Copolymers ............................................................ 12 1.2.3.2 Nanocapsules and Vesicles .................................................................... 18 1.2.3.3 Micelles.................................................................................................. 19 1.2.3.4 Nanospheres........................................................................................... 21 1.2.4 Rationale for the use of Nanoparticulate Drug Delivery Systems................ 25 1.2.4.1 Prolonged Circulation and the Enhanced Permeation and Retention Effect...................................................................................................... 25 1.2.4.2 Platforms for Active Targeting .............................................................. 28 1.2.4.3 Drug Solubilization................................................................................ 29 1.2.4.4 Evasion of Multi Drug Resistance Mechanisms.................................... 30 1.3 Polymeric Drug Delivery..................................................................................... 31 1.3.1 Polymer Molecular Weight........................................................................... 32 1.3.2 Polymer Morphology.................................................................................... 37 1.3.3 Controlled Drug Release Systems ................................................................ 42 1.3.3.1 Solvent Controlled Release.................................................................... 44 1.3.3.2 Diffusion Controlled Release................................................................. 44 1.3.3.3 Chemically Controlled Release.............................................................. 46 1.3.4 Targeted Drug Delivery ................................................................................ 49 1.3.5 Polymers for Drug Delivery.......................................................................... 49 1.3.5.1 Biodegradable Polymers for Drug Delivery .......................................... 50 1.4 Paclitaxel.............................................................................................................. 52 1.4.1 Chemistry...................................................................................................... 52 1.4.2 Pharmacology ............................................................................................... 55 1.4.3 Toxicity ......................................................................................................... 55 1.4.4 Pharmacokinetics .......................................................................................... 56 1.4.5 Paclitaxel Loaded Nanoparticulate Formulations......................................... 57 1.5 Physicochemical Properties, Drug Solubilization, and Hemocompatibility of Amphiphilic Block Copolymer Nanoparticulate Drug Delivery Systems .......... 61 1.5.1 Methods of Polymeric Nanoparticle Preparation and Drug Loading ........... 61 1.5.1.1 Direct Dissolution .................................................................................. 63  v 1.5.1.2 Film Casting or Solvent Evaporation..................................................... 63 1.5.1.3 Dialysis and Nanoprecipitation.............................................................. 64 1.5.1.4 Emulsification ........................................................................................ 64 1.5.2 Factors Influencing Physicochemical Properties and Drug Loading............ 65 1.5.2.1 Methods of Preparation.......................................................................... 65 1.5.2.2 Block Length.......................................................................................... 67 1.5.2.3 Drug and Polymer Compatibility........................................................... 72 1.5.3 Hemocompatibility of Polymer-Based Drug Delivery Systems................... 75 1.5.3.1 Coagulation ............................................................................................ 75 1.5.3.2 Complement System Activation ............................................................ 76 1.5.3.3 Strategies for Improving Hemocompatibility ........................................ 77 1.6 Thesis Rationale and Research Objectives .......................................................... 80 1.6.1 Rationale ....................................................................................................... 80 1.6.2 Research Objectives...................................................................................... 82 1.7 References............................................................................................................ 83 Chapter 2 Synthesis and Characterization of Methoxy Poly(Ethylene Glycol)- Block-Poly("-Caprolactone) Diblock Copolymers ................................. 102 2.1 Introduction........................................................................................................ 102 2.2 Experimental ...................................................................................................... 104 2.2.1 Materials ..................................................................................................... 104 2.2.2 Synthesis of Methoxy Poly(Ethylene Glycol)-Block-Poly("-Caprolactone) Diblock Copolymers .................................................................................. 105 2.2.3 Nuclear Magnetic Resonance Spectroscopy (NMR) .................................. 105 2.2.4 Gel Permeation Chromatography (GPC) .................................................... 106 2.2.5 Critical Micelle Concentration (CMC) and Pyrene Partition Equilibrium Constant (Kv)............................................................................................. 106 2.2.6 Hydrodynamic Diameter............................................................................. 107 2.2.7 Aqueous Solution Phase Behavior and Micelle Morphology of Short Block Length Methoxy Poly(Ethylene Glycol)-Block-Poly("-Caprolactone) Diblock Copolymers .................................................................................. 107 2.3 Results................................................................................................................ 108 2.3.1 Synthesis and Characterization of MePEG-b-PCL Diblock Copolymers .. 108 2.3.2 Aqueous Solution Phase Behaviour and Micelle Morphology of Short Block Length MePEG-b-PCL Diblock Copolymers............................................ 124 2.4 Discussion .......................................................................................................... 130 2.5 Conclusions........................................................................................................ 139 2.6 References.......................................................................................................... 141 Chapter 3 Solubilization of Hydrophobic Drugs by Methoxy Poly(Ethylene Glycol)- Block-Poly("-Caprolactone) Diblock Copolymer Micelles: Theoretical and Experimental Data and Correlations ............................................... 145 Chapter 4 In Vitro Human Plasma Distribution of Nanoparticulate Paclitaxel is Dependent on the Physicochemical Properties of Methoxy Poly(Ethylene Glycol)-Block-Poly("-Caprolactone) Nanoparticles............................... 146 4.1 Introduction........................................................................................................ 146 4.2 Experimental ...................................................................................................... 148 4.2.1 Materials ..................................................................................................... 148  vi 4.2.2 Formation and Characterization of MePEG-b-PCL Nanoparticles ............ 149 4.2.3 In Vitro Paclitaxel Release.......................................................................... 152 4.2.4 Blood Compatibility and In Vitro Paclitaxel Plasma Partitioning.............. 153 4.2.4.1 Blood Coagulation ............................................................................... 153 4.2.4.2 Hemolysis ............................................................................................ 153 4.2.4.3 In Vitro Paclitaxel Plasma Distribution ............................................... 155 4.3 Results................................................................................................................ 156 4.3.1 Formation and Characterization of MePEG-b-PCL Nanoparticles ............ 156 4.3.2 In Vitro Paclitaxel Release.......................................................................... 161 4.3.3 Blood Compatibility and In Vitro Paclitaxel Plasma Distribution ............. 164 4.3.3.1 Blood Coagulation ............................................................................... 164 4.3.3.2 Hemolysis ............................................................................................ 164 4.3.3.3 In Vitro PTX Plasma Distribution ....................................................... 167 4.4 Discussion .......................................................................................................... 173 4.5 Conclusions........................................................................................................ 181 4.6 References.......................................................................................................... 183 Chapter 5 Summarizing Discussion, Conclusions and Suggestions for Future Work ..................................................................................................................... 188 5.1 Summarizing Discussion and Conclusions........................................................ 188 5.2 Suggestions for Future Work ............................................................................. 197 5.3 References.......................................................................................................... 199 Appendix 1  Ethics Certificate ..................................................................................... 202   vii LIST OF TABLES  Table 1.1 Examples of amphiphilic block copolymers and types of nanoparticulate delivery systems........................................................................................ 16 Table 1.2 Nanoparticulate drug delivery systems formed by amphiphilic block copolymers and their general characteristics ............................................ 24 Table 2.1 Synthesis and characterization data for methoxy poly(ethylene glycol)- block-poly("-caprolactone) diblock copolymers .................................... 111   viii  LIST OF FIGURES  Figure 1.1 Classification of nanoparticles based on constitutive materials. Inside circle indicates materials used to form nanoparticles. Outside circle includes examples of each classification of nanoparticle. Intersecting lines indicate nanoparticles composed of a blend of materials. .......................... 7 Figure 1.2 Polymer classifications based on structure. Cubes and spheres represent different monomers. .................................................................................. 14 Figure 1.3 Representation of a molecular weight distribution curve with the various average molecular weights indicated. ....................................................... 34 Figure 1.4 Schematic representation of the fringed micelle model of polymer crystallinity. .............................................................................................. 39 Figure 1.5 Models of reentry of chains for the chain folded model of polymer crystallinity. A) regular reentry, B) irregular reentry and C) switchboard reentry. ...................................................................................................... 41 Figure 1.6 Schematic of controlled drug release systems. A) Reservoir system and B) matrix system. Arrows represent transport of drug out of the device....... 43 Figure 1.7 Schematic of bulk and surface degradation mechanisms of polymer degradation during controlled drug release............................................... 48 Figure 1.8 The chemical structure of paclitaxel. ........................................................ 54 Figure 1.9 Methods of micelle preparation and drug loading. A) direct dissolution, B) nanopreciptation or dialysis, C) film casting or solvent evaporation D) emulsification............................................................................................ 62 Figure 2.1 Polymerization reaction of MePEG-b-PCL. ........................................... 110 Figure 2.2 GPC chromatograms of the polymerization mixture with a feed ratio of 60:40 of MePEG 750 g/mol and CL. Reaction times are indicated on the left. At 0 hrs, peak A represents MePEG and peak B is CL. .................. 112 Figure 2.3 The effect of reaction time on the number average molecular weight (Mn) (!) of MePEG-b-PCL and weight percent of unreacted caprolactone during the synthesis of MePEG17-b-PCL4 ("). Data points and error bars represent the mean ±  S.D. (n =3)............................................................ 113 Figure 2.4 A representative 1 H NMR spectrum of a MePEG-b-PCL diblock copolymer using a Bruker 400 MHz spectrometer. ................................ 115 Figure 2.5 Fluorescence excitation spectra of pyrene solubilized in aqueous solutions of MePEG17-b-PCL4. Emission wavelength 390 nm and 37°C. ............ 118 Figure 2.6 Representative plot of I336nm/I333nm as a function of copolymer concentration for the determination of CMC and Kv. ............................ 121  ix Figure 2.7 Plots of pyrene fluorescence intensity ratio I336nm/I333nm at 37°C as a function of polymer concentration for the determination of CMC. MePEG17-b-PCL2 (!),MePEG17-b-PCL3 (!),MePEG17-b-PCL4 (!), MePEG17-b-PCL10 ("), MePEG44-b-PCL7 (#), MePEG44-b-PCL12 (!) and MePEG114-b-PCL19 ("). Data points and error bars represent the mean ± S.D. (n =3). ............................................................................................. 122 Figure 2.8 A representative plot of (F-Fmin)/Fmax-F) vs. polymer concentration for the determination of the pyrene partition coefficient (Kv) for the copolymer MePEG17-b-PCL10. ................................................................................. 123 Figure 2.9 Photographs of 20 g/L aqueous solutions of MePEG17-b-PCL2 at 4°C (below the Krafft point), 37°C (micellar solution), and 60°C (above the cloud point). ............................................................................................ 126 Figure 2.10 Phase diagrams of MePEG-b-PCL diblock copolymers with a MePEG MW of 750 g/mol. (A) MePEG17-b-PCL10, (B) MePEG17-b-PCL4, (C) MePEG17-b-PCL2. (!) Cloud point, (") Krafft point, and (#) CMC. Data points and error bars represent the mean ±  S.D. (n =3).......................... 127 Figure 2.11 Transmission electron micrographs of MePEG-b-PCL micelles. Shadowed at a 30°  angle with Pt/Pd alloy. Accelerating voltage 80kV. (A) MePEG17- b-PCL2 (B) MePEG17-b-PCL4, (C) MePEG17-b-PCL10.......................... 129 Figure 2.12 Schematic of the concentration (dC) and temperature (dT) dependent aqueous phase behavior of short block length MePEG-b-PCL diblock copolymers. ............................................................................................. 138 Figure 4.1 Dependence of nanoparticle hydrodynamic diameter on copolymer concentration for MePEG114-b-PCL19 micelles (!) and MePEG114-b- PCL104 nanospheres ("). All points represent the mean hydrodynamic diameter ±  SD (n = 3). ............................................................................ 158 Figure 4.2 Cryo TEM images of (A) MePEG114-b-PCL19 micelles and (B) MePEG114- b-PCL104 nanospheres. Bar represents 100 nm. ...................................... 159 Figure 4.3 Anisotropy as a function of temperature. Spingomyelin bilayer (!) MePEG114-b-PCL19 micelles (#) MePEG114-b-PCL104 nanospheres ($) polysorbate 20 micelles ("). All points represent the mean anisotropy ± SD (n = 3). .............................................................................................. 160 Figure 4.4 Solubilization of PTX by (A) MePEG114-b-PCL19 micelles and (B) MePEG114-b-PCL104 nanospheres formed by nanoprecipitation of copolymer and drug in DMF solutions. All nanoparticles were formed in PBS (0.01M, pH 7.4) using copolymer concentrations of 0.36mM. [PTX] solubilized (%), loading efficiency (!). Data points and error bars represent the mean ±  SD (n=3)............................................................... 162 Figure 4.5 Release of PTX from dialysis bags. Drug was either in its free form (!) or solubilized in MePEG114-b-PCL19 micelles (") or MePEG114-b-PCL104  x nanospheres (#) and released into PBS (0.01M, pH 7.4) at 37°C. Data points and error bars represent the mean ±  SD (n=3)............................. 163 Figure 4.6 Effect of nanoparticle type and concentration on the (A) prothrombin time (PT) and (B) activated partial thromboplastin time (APTT). White bars represents control (PBS), black bars represents MePEG114-b-PCL19 micelles, grey bars represents MePEG114-b-PCL104 nanospheres. All bars represent mean clotting time ±  SD (n=3). * p <0.05, different from control. .................................................................................................... 165 Figure 4.7 Time course of hemolysis by MePEG17-b-PCL4 micelles (!), MePEG114- b-PCL19 micelles ("), and MePEG114-b-PCL104 nanospheres (#) at 37°C at copolymer concentrations of 0.1% w/v. Data represents mean ±  SD (n = 3). ............................................................................................................ 166 Figure 4.8 Partitioning of PTX into various density fractions after a 1 hr incubation in distilled water. Drug was formulated in either MePEG114-b-PCL19 micelles (white) or MePEG114-b-PCL104 nanospheres (black). All values represent mean ±  SD (n=6). Inset is a schematic of an ultracentrifuge tube with the position of the various density fractions and their corresponding lipoprotein fractions after ultracentrifugation. NP fraction represents the clearly separated white band of MePEG114-b-PCL104 nanospheres. ....... 169 Figure 4.9 Time dependence of partitioning of free PTX into various human plasma fractions during incubation times of 1 hr (black), 6 hrs (white), 12 hrs (grey), and 24hrs (diagonal stripes). All values represent mean ±  SD (n=6). * p < 0.05, different from all other samples in plasma fraction... 170 Figure 4.10 Time dependence of partitioning of MePEG114-b-PCL19 micellar PTX into various human plasma fractions during incubation times of 1 hr (black), 6 hrs (white), 12 hrs (grey), and 24hrs (diagonal stripes). All values represent mean ±  SD (n=6) All values represent mean ±  SD (n=6). # p < 0.05, different from all other samples in plasma fraction. * p < 0.05, different from 1 hr sample in plasma fraction. ....................................... 171 Figure 4.11 Time dependence of partitioning of MePEG114-b-PCL104 nanosphere PTX into various human plasma fractions during incubation times of 1 hr (black), 6 hrs (white), 12 hrs (grey), and 24hrs (diagonal stripes). Hatched bars represent distribution of PTX after incubation of free PTX for 10 min followed by addition of blank MePEG114-b-PCL104 nanospheres and an additional 1 hr incubation. All values represent mean ± SD (n=6). * p < 0.05, different from all other samples in plasma fraction. # p < 0.05, different from 1 hr sample in plasma fraction. ....................................... 172   xi  LIST OF ABBREVIATIONS #   Hildebrand solubility parameter #d   Hildebrand solubility parameter (dispersion components) #h   Hildebrand solubility parameter (hydrogen bonding components) #p   Hildebrand solubility parameter (polar components) #t    Total Hildebrand solubility parameter $   Density of PCL in the core %sp   Flory Huggins interaction parameter Apo-E   Apolipoprotein E APTT   Activated partial thromboplastin time ATP   Adenosine triphosphate C   Copolymer concentration C0   Total concentration of drug in the matrix (dissolved and dispersed) CL   "-caprolactone CMC   Critical micelle concentration CNS   Central nervous system cryo TEM  Cryogenic transmission electron microscopy Cs(m)   Solubility of the drug in the polymer matrix CWC   critical water content D   Diffusion coefficient DLS   Dynamic light scattering DMF   N,N,-dimethyl formamide  xii DMSO   Dimethyl sulfoxide DPH   1,6-diphenyl-1,3,5-hexatriene Ecoh   Cohesive energy ecoh   Cohesive energy density Ed    Dispersion forces Eh    Hydrogen bonding EM   Electron microscope Ep    Polar cohesive forces EPR   Enhanced permeability and retention GPC   Gel permeation chromatography HDL   High density lipoprotein HPLC   High performance liquid chromatography i.p.   Intraperitoneal i.v.   Intravenously IgG   Immunoglobulin G Kv   Pyrene partition equilibrium coefficient l   Thickness of the device LDL   Low density lipoprotein LPDP   Lipoprotein deficient plasma fraction ! Mk    Average molecular weight ! Mn    Number averaged molecular weight ! M"    Viscosity averaged molecular weight ! Mw    Weight averaged molecular weight  xiii ! Mz   Z averaged molecular weight M0    Initial amount of drug loaded into the device MAC   Membrane attack complex MePEG  Methoxy poly(ethylene glycol) MPS   Mononuclear phagocyte system Mt    Amount of drug remaining in the device at any given time t MTD   Maximum tolerated doses MW   Molecular weight Nagg   Aggregation number “NP fraction”  Nanoparticle fraction NMR   Nuclear magnetic Resonance Spectroscopy PAA   Poly(acrylic acid) PBS   Phosphate buffered saline PCL   Poly("-caprolactone) PDI   Polydispersity index PDLLA  Poly(d,l lactic acid) PEG   Poly(ethylene-glycol) PEO   Poly(ethylene oxide) P-gp   P-glycoprotein PLA   Poly(lactic acid) PLGA   Poly(d,l-lactic-co-glycolic acid) PMMA  Poly(methyl methacrylate) PS   Polystyrene  xiv PT   Prothrombin time PTX   Paclitaxel PVP   Poly(N-vinyl-2-pyrrolidone) QDs   Quantum dots r   Fluorescence anisotropy R   Gas constant RES   Reticuloendothelial system Rhyd    Hydrodynamic radius ro   Radius of the sphere SMANCS  Neocarzinostatin conjugated to copoly(styrene-maleic-acid) T   Absolute temperature t1/2&   Distribution half life t1/2'   Elimination half life TEM   Transmission electron microscopy THF   Tetrahydrofuran TPGS   d-&-tocopheryl polyethylene glycol succinate UV VIS  Ultraviolet – visible spectroscopy V   Volume VLDL   Very low density lipoprotein / chylomicrons Vs    Molar volume of the solute  xv  ACKNOWLEDGEMENTS This thesis would not have been possible without the support of many people. I would like to, first and foremost, thank my supervisor, Dr. Helen Burt. Her knowledge, dedication, enthusiasm and support have been instrumental in my personal and professional development. I could not have asked for a better mentor. Thank you to John Jackson, who was always been there to provide technical expertise and insightful scientific discussion.  My sincere thanks goes to my co supervisor Dr. Richard Liggins and my committee members: Drs. Marcel Bally, Colin Fyfe, Brian Rodrigues, and Kishor Wasan for their effort and thoughtful discussions. I am grateful for the technical expertise of Dr. Rajesh Kainthan and Samuel Hester for help with the plasma coagulation experiments, Olena Sivak, Steve Lee and Ali Reza Merali for help with the plasma partitioning experiments. Thank you to Karen Chu and Framin Mark for their hard work through the summers. Thank you to all my past, present and honorary lab mates; Jason Zastre, Chris Springate, Tobi Higo, Karen Long, Linda Liang, Ruiwen Shi, Chiming Yang, Sam Gilchrist, Antonia Tsallas, John Lu, Rita Zhao, Praveen Elamanchili, Clement Mugabe, Melanie ter Borg, Katherine Haxton and Linda Tran, for your support, suggestions and most of all, making the Burt Lab a great place to be.  I am truly grateful for the support and encouragement of my parents over the years. To my brother Ryan and friend Tom, for always providing an avenue for stress relief, whether it was on foot, bicycle or skis, thank you, it was much needed and appreciated.  xvi  Financial support from the Canadian Institutes of Health Research and University Graduate Fellowship is gratefully acknowledged.  xvii CO-AUTHORSHIP STATEMENT This thesis is comprised of three manuscripts of which I am the principal author. In all cases, I was the primary individual responsible for the identification and design of the research program, practical aspects of the research, data analyses and manuscript preparation. The contribution of co-authors was through the provision of intellectual discussion.   1  CHAPTER 1 INTRODUCTION 1.1 PROJECT OVERVIEW Amphiphilic block copolymers are composed of a hydrophobic block covalently bound to a hydrophilic block creating a polymeric material that has opposite affinities for aqueous solvents. In aqueous milieu these materials are capable of forming nano-sized structures possessing a core-shell morphology, which is characterized by a hydrophobic core surrounded by a highly water bound, hydrophilic corona. These nanoparticles have great potential as drug delivery systems due to their ability to increase the aqueous solubility of hydrophobic drugs and increase their blood circulation time.  Typically, the hydrophilic block is poly(ethylene-glycol) (PEG) or methoxy poly(ethylene glycol) (MePEG). A PEG surface coating acts to reduce the adsorption of plasma proteins such as opsonins and apolipoproteins on the surface of nanoparticles. This prevents the recognition of the particles by the mononuclear phagocytic system, prolonging the circulation of PEGylated nanoparticles in blood, potentially resulting in the passive targeting of disease sites (Gref et al., 2000). PEG coated nanoparticles can also resist adsorption of proteins of the coagulation cascade, thus, preventing thrombosis and increasing the biocompatibility of these systems (Sahli et al., 1997). These shielding effects require sufficient PEG chain length and density at the surface of the particles to effectively repel plasma proteins (Gref et al., 2000).  A variety of core forming blocks have been evaluated; however, the most common choices are biodegradable and biocompatible polyesters such as poly(d,l lactic acid)  2 (PDLLA) (Zhang et al., 1996), poly(caprolactone) (PCL) (Allen et al., 2000) or poly(lactic-co-glycolic acid) (PLGA) (Yoo and Park, 2001). The hydrophobic core has been utilized as a cargo space for poorly water-soluble drugs, increasing their aqueous solubility and providing controlled release. The extent of micellar solubilization of hydrophobic solutes varies depending on the nature of the core-forming block. It has been suggested that the reason for this preferential solubilization is due to the compatibility between the polymer and solute as estimated by the Flory-Huggins interaction parameter (%sp) (Nagarajan, 2001; Nagarajan et al., 1986; Nagarajan and Ganesh, 1996). This theory has been used for qualitative comparisons of the extent of solubilization of different drugs (Allen et al., 1999; Lavasanifar et al., 2002; Liu et al., 2004; Soo et al., 2002),. However, few, if any, studies have systematically quantified the micellar solubilization as a function of the %sp or the core-forming block length of the copolymer.  Amphiphilic block copolymer nanoparticles can be categorized in terms of structure, including micelles, nanospheres, crew-cut micelles, and polymersomes. The type of structure formed is dependent on the relative lengths of the hydrophobic and hydrophilic blocks of the copolymer, molecular weight, as well as the method of preparation (Letchford and Burt, 2007). Although micelles and nanospheres both have a core-shell architecture, they differ in a few fundamental ways. The copolymer chains of micelles formed from water-soluble amphiphilic copolymers are in a dynamic equilibrium with free unimers in solution (Alexandridis and Hatton, 1995) and therefore should not be considered solid particles, but rather association colloids (Martin, 1993). As the length of the hydrophobic block increases, the copolymer becomes water insoluble. In order to  3 form nanoparticles from these relatively hydrophobic copolymers, they must first be solubilized in an organic solvent prior to emulsification or precipitation. The resulting structures exhibit a core-shell architecture, similar to micelles but are generally larger and possess a more solid-like core and are considered to be a separate phase (Gref et al., 1994). These nanoparticles are termed nanospheres. It has been demonstrated that there are differences in the physicochemical properties of micelles and nanospheres, such as mobility of the core blocks and thermodynamic and kinetic stabilities; however, little is known about how these two types of nanoparticles interact with blood and plasma constituents or how the encapsulated drug distributes into plasma fractions.  Paclitaxel (PTX) is a potent, hydrophobic anticancer drug that is commercially formulated in a mixture of Cremophor( EL and ethanol. This excipient has been associated with a number of serious adverse effects (Dorr, 1994); hence, there has been considerable interest in the development of Cremophor( EL-free paclitaxel formulations. Previously, our group developed a MePEG-b-PDLLA micellar formulation of PTX (Zhang et al., 1996). It was found that upon incubation with human plasma, both free and micellized PTX rapidly distributed into lipoprotein and lipoprotein deficient plasma fractions suggesting that the drug was released from the micelles and associated with the plasma proteins (Ramaswamy et al., 1997). Despite numerous pharmacokinetic studies of non-drug loaded diblock copolymer micelles showing that copolymer nanoparticles are capable of prolonged circulation in the blood (Liu et al., 2007; Yamamoto Y, 2001), it is clear that for some PTX loaded nanoparticles, PTX does not remain with the carrier (Burt et al., 1999; Kim et al., 2001; Le Garrec et al., 2004). It is evident that in order to  4 increase the blood circulation time of PTX, better drug retention within the nanoparticle core is necessary. Thus, one possible strategy is to solubilize the drug in the more solid- like core of a nanosphere. We hypothesized that nanospheres, characterized by an increased thermodynamic and kinetic stability as compared to micelles, may prevent the rapid dissociation of PTX from the carrier.  The overall objective of this project was to characterize a series of MePEG-b-PCL diblock copolymers and correlate the physicochemical properties of these systems with their performance as nanoparticulate drug delivery systems. The first part of the thesis focuses on synthesis and characterization of a series of MePEG-b-PCL diblock copolymers with varying hydrophilic and hydrophobic block lengths. The physicochemical properties of the micelles formed by these materials are assessed and related to systematic changes in the hydrophilic and hydrophobic block lengths. The second part of the thesis investigates the development of empirical relationships between the extent of solubilization of hydrophobic drugs and the compatibility between the core - forming block and the drug as well as the length of the core-forming block. The final chapter compares micellar and nanosphere formulations composed of MePEG-b-PCL, their hemocompatibility and differences in the in vitro plasma distribution of PTX, solubilized by these two nanoparticulate delivery systems. 1.2 NANOPARTICULATE DRUG DELIVERY SYSTEMS 1.2.1 Definitions and Classification Nanotechnology may be defined as “the creation of functional materials devices and systems through control of matter on the nanometer scale and the exploitation of novel  5 properties and phenomenon developed” (http://www.lanl.gov/mst/nano/definition.html). This field has had an impact on virtually every industry and is particularly evident in the computer industry with the rapid development and miniaturization of computer chips. It is estimated that the United States will spend an estimated 3.7 billion dollars on nanoscience and nanotechnology between 2005 and 2008. Furthermore, it has been stated that nanotechnology could have as much of an impact as inventions such as the steam engine, electricity or the internet (Hassan, 2005). One of the greatest contributions of nanotechnology to human kind may be through application of these technologies to healthcare, a field termed nanomedicine. Nanomedicine requires the blending of multiple disciplines such as biology, chemistry, physics, chemical and mechanical engineering, material science and clinical medicine (Farokhzad and Langer, 2006) and is poised to change the diagnosis, treatment and prevention of disease in the future. Some of the first applications of nanotechnology to medicine were the use of nanoparticles.  In recent years there has been significant interest in the use of nanoparticles for drug delivery applications. For the purpose of drug delivery, nanoparticles are colloidal-sized particles, possessing diameters ranging between 1-1000 nm, and a large number of drugs have been loaded into these nanoparticles using a variety of different methods. Nanoparticles composed of a range of materials including lipids, polymers, and inorganic materials have been developed, resulting in delivery systems that vary in their physicochemical properties and thus their applications (Gabizon et al., 2003; Klumpp et al., 2006; Liggins and Burt, 2002; Paciotti et al., 2006). One of the earliest types of nanoparticles were lipid vesicles known as liposomes, which were described as early as 1964 (Bangham and Horne, 1964). Liposomes have since been utilized as drug delivery  6 systems and commercialized products for drugs such as doxorubicin and amphotericin B. To date, an array of nanoparticulate drug delivery systems have been investigated including, but not limited to, liposomes, micelles, nanospheres, niosomes, nanocapsules, solid lipid nanoparticles, microemulsions and carbon nanotubes and may be broadly classified according to the materials from which they are composed (figure 1.1).  7         Figure 1.1 Classification of nanoparticles based on constitutive materials. Inside circle indicates materials used to form nanoparticles. Outside circle includes examples of each classification of nanoparticle. Intersecting lines indicate nanoparticles composed of a blend of materials.   8 1.2.2 Therapeutic Applications Nanoparticles find application in a wide array of medical fields; however, the most products either in development or on the market are drug delivery systems (Wagner et al., 2006). As of 2004 there were 23 nanotechnology based therapeutic drug delivery products on the market for indications including cancer, fungal infection and hepatitis C (Zhang et al., 2008). Most of these products are based on either liposomal delivery systems or polymer/drug conjugates.  Anticancer drug loaded nanoparticulate formulations have been investigated for decades. It was established that the small size of these drug carriers prolonged the circulation of the entrapped drug leading to passive targeting of tumors when administered intravenously, particularly if they were surface coated with poly(ethylene glycol) (Torchilin, 2005). The liposomal doxorubicin formulations, Myocet( and Doxil( were the first to be approved with indications for metastatic breast cancer and ovarian cancer (Zhang et al., 2008). Recently, albumin bound paclitaxel nanoparticles (Abraxane() have been approved for treatment of metastatic breast cancer (Gradishar, 2006). Significant ongoing research in the field involves active targeting and controlled or triggered release of drug payloads to enhance efficacy and reduce toxicity (Mahmud et al., 2007). Targeting has been achieved by the conjugation of targeting antibodies (Schnyder et al., 2005), peptides (Schiffelers et al., 2003), sugars (Nagasaki et al., 2001) and aptamers (Farokhzad et al., 2006) to the surface of nanoparticles and has been met with some success. These molecules bind to specific antigens or receptors that are overexpressed by the targeted cell type allowing for the selective targeting of disease sites. Once a  9 nanoparticle reaches its intended site, drug release may be triggered through various means, including externally applied stimuli such as externally applied ultrasound or heat. Ultrasound acts to rupture the carrier, releasing the drug payload and/or increase the permeability of the surrounding cells (Pitt et al., 2004). Externally applied heat may be used to trigger the release of drug from nanoparticles composed of thermally sensitive copolymers of poly(N-isopropylacrylamide) through a thermally induced phase transition of the polymer (Kohori et al., 1998).  Nanoparticles possess significant potential as a platform for targeting molecular imaging labels for diagnostic purposes. Of particular importance are nanoparticles formed from inorganic fluorescent semiconductors, termed quantum dots (QDs). Due to their high quantum yield, high molar extinction coefficients, strong resistance to photobleaching, narrow emission spectra and size dependent emission wavelength, QDs are particularly well suited for biomedical imaging and consistently outperform traditional organic dyes (Cai et al., 2007). QDs have been extensively used in in vitro and cell based applications such as cellular labeling and cell tracking. Recently, in vivo applications of QDs conjugated with targeting ligands such as peptides and antibodies, have shown promise in imaging specific tissues including neovasculature, tumor cells and lymphatic vessels of breast and prostate cancer tumors (Akerman et al., 2002; Gao et al., 2004). Other imaging modalities are also utilizing the benefits of nanoparticles. Encapsulation of contrast agents in nanoparticles and targeting to tissues of interest has been a topic of active research for imaging techniques including targeted ultrasound, computed tomography, and magnetic resonance imaging (Cai and Chen, 2007).  10  Many of the products of drug discovery are poorly water-soluble compounds and protein and peptide molecules, which unfortunately do not lend themselves to oral delivery, the preferred method of drug administration. The development of nano-sized formulations has shown promise in increasing the water solubility of hydrophobic drugs. Such formulations include NanoCrystal( technology and the production of nanosuspensions (Merisko-Liversidge et al., 2003; Muller et al., 2001). The increased surface area of the drug particles results in an increased dissolution rate and improved bioavailability. The use of nanoparticles for the delivery of biomolecules has two advantages. Firstly, the encapsulation of proteins and peptides in particulate carriers aids in the protection of the therapeutic against degradation by acid and enzymes. Secondly, these particles have been demonstrated to be subject to translocation across the intestinal mucosa by various mechanisms leading to improved oral bioavailiablity (Jung et al., 2000). Several sites and mechanisms of transport of particles across the intestinal wall exist including paracellular and endocytotic pathways as well as the gut-associated lymphoid tissue (Jung et al., 2000). Oral nanoparticulate delivery systems for a number of protein and peptide drugs have been investigated including insulin, calcitonin and leutinizing hormone releasing hormone (Allemann et al., 1998).  It is estimated that 1.5 billion people suffer from central nervous system (CNS) disorders worldwide including such diseases as Alzheimer’s disease and cerebral ischemia (Tiwari and Amiji, 2006) leading to an increase in the development of CNS therapeutics. However, treatment of these conditions has been hampered due to the inability to  11 efficiently deliver therapeutic agents to the brain. There are several physical barriers that restrict the movement of drugs into the CNS including endothelial tight junctions, the presence of efflux transporters and a deficiency of pinocytic vesicles (Tiwari and Amiji, 2006). Effective brain accumulation of drugs encapsulated in nanoparticles such as immunoliposomes and surfactant coated poly(butyl)cyanoacrylate nanoparticles has been demonstrated, but the exact mechanism of nanoparticulate uptake into the CNS is not clear (Emerich and Thanos, 2007). It has been suggested that in situ coating of nanoparticles with apolipoprotein-E and subsequent uptake by the low density lipoprotein uptake system, as well as inhibition of efflux transporters by surfactant coatings and endocytosis by capillary endothelial cells may play key roles (Tiwari and Amiji, 2006).  As posterior segment ocular diseases including age related macular degeneration and diabetic retinopathy become more prevalent in industrialized countries, there is a need to effectively deliver drugs to these disease sites (del Amo and Urtti, 2008). Since adequate drug concentrations are not achievable in the posterior tissues by local administration, the only viable means of drug delivery to these tissues are by intravitreal injection or systemic administration. Unfortunately, both of these methods have their disadvantages. Due to the invasive nature of intravitreal injections, it is necessary that the drugs have prolonged retention in the eye to minimize the number of injections. Studies have shown that intravitreal injection of nanoparticles results in localization in the retinal pigment epithelium for prolonged periods of time thus providing sustained delivery of drugs (Bejjani et al., 2005; Bourges et al., 2003). These therapies are being investigated for the treatment of macular degeneration and diabetic retinopathy (Mo et al., 2007; Xu et al.,  12 2007). As only a small amount of blood flows through the posterior ocular segment, in order for systemic circulation to be a feasible delivery route, targeting to the cells in this tissue including the retinal pigment epithelium, choroid and retinal capillaries is desirable. Intravenously injected therapies that are targeted to specific tissues of the eye using peptides and antibodies are currently being investigated for photodynamic therapy of macular degeneration (Birchler et al., 1999; Renno et al., 2004).  1.2.3 Amphiphilic Block Copolymer Drug Delivery Systems Nanoparticulate drug delivery systems have been developed using a broad range of different homopolymers and copolymers. Amphiphilic block copolymers have proved to be particularly interesting in terms of their ability to form nanoparticulates.  1.2.3.1 Amphiphilic Block Copolymers A homopolymer is a polymer that is made of only one type of monomer. Examples of homopolymers are poly(lactic acid) and poly(glycolic acid). If the polymer is made of two or more types of repeat units, the structure is referred to as a copolymer. Copolymers may be subdivided into four main categories (Figure 1.2):  (1) Random copolymers where the monomers repeat in a random order, as the name suggests. (2) Alternating copolymers have a perfectly alternating arrangement of monomers along the chain.  13 (3) Block copolymers contain substantial sequences or blocks of each monomer. These polymers typically are diblock or triblock depending on the number of blocks. (4) Graft copolymers contain one main chain of monomers with a series of a different monomer branching from this structure.  Network polymers are formed when branches from the main polymer chains are bonded or cross-linked to other polymer chains. An example of a network polymer is epoxy resin.  14                       Figure 1.2 Polymer classifications based on structure. Cubes and spheres represent different monomers.   15  The use of amphiphilic block copolymers of drug delivery applications has been increasingly more popular over the past two decades. These materials are composed of two different monomers, one hydrophilic and the other hydrophobic. The monomers are arranged so that the polymer has significant sections, or blocks, entirely composed of one type of monomer. Copolymers are typically arranged in either diblock or triblock configurations. Since the different copolymer blocks have opposite affinities for an aqueous phase, copolymers will orient themselves to minimize the interaction between the hydrophobic blocks and the aqueous phase. In aqueous milieu, amphiphilic copolymers may self assemble to form aggregates termed micelles, if they possess the appropriate proportion of hydrophilic and hydrophobic block lengths. Micelles are nano- sized particles characterized by a core consisting of hydrophobic blocks surrounded by a corona composed of highly water bound hydrophilic blocks. The use of micelles as drug delivery systems is well documented in the literature and will be described in more detail below. Amphiphilic block copolymers have been investigated in a number of other nanoparticulate drug delivery systems such as nanospheres (Gref et al., 1994), polymersomes (Ahmed and Discher, 2004), and nanocapsules (de Faria et al., 2005) (see Table 1.1), as well as other delivery systems including thermoreversible gels(Jeong et al., 2000), injectable pastes (Jackson et al., 2000) and  as excipients in solid matrices for the modulation of drug release (Jackson et al., 2007).  16 Table 1.1 Examples of amphiphilic block copolymers and types of nanoparticulate delivery systems.   Name and Structure Abbreviation Delivery Systems References Methoxy poly(ethylene glycol)-b-poly(caprolatone) CH3O CH2CH2O n C(CH2)5O O m H  MePEG-b-PCL Micelles Nanospheres Nanocapsules Polymersomes (Allen et al., 2000) (Kim et al., 2001) (Ameller et al., 2003) (Ahmed et al., 2004) Methoxy poly(ethylene glycol)-b-poly(d,l lactide) CH3O CH2CH2O n CCHO O m H CH3 MePEG-b-PDLLA Micelles Nanospheres Nanocapsules Polymersomes (Burt et al., 1999) (Ameller et al., 2003) (de Faria et al., 2005) (Meng et al., 2003) Poly(ethylene oxide)-b-poly(!-benzyl L aspartate) HO CH2CH2O n CCHNH O m H CH2 CO O CH2  PEO-b-PBLA Micelles (Kataoka et al., 2000) Poly(ethylene oxide)-b-poly(propylene oxide)-b- poly(ethylene oxide) or Pluronic" HO CH2CH2O n CHCH2O m n CH3 CH2CH2O OH  PEO-b-PPO-b-PEO Micelles (Alexandridis et al., 1994)  17 In the majority of amphiphilic copolymers, the hydrophilic block consists of poly(ethylene glycol) (PEG), also known as poly(ethylene oxide) (PEO). PEG is chosen due to its water solubility and high degree of biocompatibility. Often the PEG used is terminated with !-methoxy in order to make one end of the polymer inert so that the "- hydroxyl end of the PEG can be used as an initiator for polymerization of hydrophobic monomers. There has been growing interest in synthesizing PEGs that bear groups that are not inert, such as amino (NH2-PEG-OH) and carboxyl (COOH-PEG-OH) end groups. With these variations it is possible to chemically bond targeting moieties to PEG for the preparation of targeted nanoparticles. Hydrophilic blocks other than PEG such as poly(N- vinyl-2-pyrrolidone) (PVP) have been used as the hydrophilic block in PVP-b-PDLLA copolymers by Le Garrec et al. (Le Garrec et al., 2004). It has been shown that PVP in its hydrated state has some capacity to solubilize drugs leading to a greater drug loading capacity of PVP based micelles as compared to PEG based micelles (Benahmed et al., 2001). Other hydrophilic blocks such as poly(acrylic acid) have been investigated in order to prepare mucoadhesive delivery systems (Inoue et al., 1998).  A number of hydrophobic blocks have been investigated for the preparation of amphiphilic block copolymers, the most common of which are the polyesters. Copolymers with hydrophobic blocks of poly(d,l lactic acid) (PDLLA) (Ha and Kim, 1999), poly(glycolic-co-lactic acid) (PLGA) (Yoo and Park, 2001) and poly(caprolactone) (PCL) (Allen et al., 2000) have been extensively investigated. The popularity of these polymers is due to their known biocompatibility as evidenced by their approval by the FDA and an extensive history of use in biomedical applications.  18 Furthermore, these materials provide a range of degradation times providing a degree of control over the release rate of the encapsulated drug. Another popular choice for hydrophobic block is the poly amino acids, including poly(aspartic acid) and poly(!- benzyl-L-aspartate). Successful formulations of doxorubicin were prepared by conjugating the drug to these blocks (Kwon et al., 1994; Yokoyama et al., 1991). Commercially available triblock copolymers called Pluronics", composed of PEO hydrophilic blocks and poly(propylene oxide) hydrophobic blocks, have been the subject of many studies investigating their thermoreversible nature (Strappe et al., 2005) as well as their potential as micellar delivery systems (Rapoport, 1999) and modulators of drug efflux proteins (Batrakova et al., 1998; Kabanov et al., 2002). 1.2.3.2 Nanocapsules and Vesicles Polymeric nanocapsules and polymersomes are colloidal sized, vesicular systems in which the drug is confined to a reservoir or within a cavity surrounded by a polymer membrane or coating (Soppimath et al., 2001) (Table 1.2). Frequently, if the core is an oily liquid and the surrounding polymer is a single layer of polymer, the vesicle is referred to as a nanocapsule. These systems have found utility in the encapsulation and delivery of hydrophobic drugs including Ru 58668, methotrexate, xanthone and 3- methylxanthone (Ameller et al., 2003; de Faria et al., 2005; Teixeira et al., 2005). Polymers used for the formation of nanocapsules have typically included polyester homopolymers such as PDLLA, PLGA and PCL. Copolymers of PEG and PDLLA have been used to avoid opsonization of the particles, similar to nanospheres (Ameller et al., 2003). Nanocapsules composed of a copolymer of PEG and chitosan have been used for the oral delivery of salmon calcitonin. The PEG was found to increase the stability the  19 nanocapsules in gastrointestinal fluid while reducing their cytotoxicity (Prego et al., 2006).  If the core of the vesicle is an aqueous phase and the surrounding coating is a polymer bilayer, the particle is referred to as a polymersome (Table 1.2) (Discher and Eisenberg, 2002). These vesicles are analogous to liposomes and find utility in the encapsulation and delivery of water-soluble drugs, which can be entrapped in their aqueous reservoir, but they differ from liposomes in that the external bilayer is composed of amphiphilic copolymers. The diblock copolymers composed of PEG with either poly(butadiene) or poly(ethylethylene) are strong vesicle or polymersome formers (Bermudez et al., 2002; Discher et al., 1999). These materials are bioinert but not biodegradable, and therefore investigations have focused on the development of polymersomes composed of PEGylated polyesters such as PEG-b-PDLLA and PEG-b-PCL either as the sole constituent of the vesicle or blended with PEG-b-poly(butadiadiene) (Ahmed and Discher, 2004; Meng et al., 2003). Polymersomes generally possess a greater PEG surface density and longer circulation times compared to PEGylated liposomes (Photos et al., 2003). 1.2.3.3 Micelles Due to the unique structure of amphiphilic molecules they have a tendency to accumulate at the boundary of two phases and thus are termed surfactants. In aqueous solutions, amphiphilic molecules orientate themselves so that the hydrophobic blocks are removed from the aqueous environment in order to achieve a state of minimum free energy. As the concentration of amphiphile in solution is increased, the free energy of the system begins  20 to rise due to unfavourable interactions between water molecules and the hydrophobic region of the amphiphile resulting in structuring of the surrounding water and a subsequent decrease in entropy. At a specific and narrow concentration range of amphiphile in solution, termed the critical micelle concentration (CMC), several amphiphiles will self-assemble into colloidal sized particles termed micelles (Table 1.2). The formation of micelles effectively removes the hydrophobic portion of the amphiphile from solution minimizing unfavorable interactions between the surrounding water molecules and the hydrophobic groups of the amphiphile. If the amphiphile concentration in solution remains above the CMC, micelles are thermodynamically stabilized against disassembly. Upon dilution below the CMC, micelles will disassemble, the rate of disassembly being largely dependent on the structure of the amphiphiles and interactions between the chains (Allen et al., 1999). In this respect, amphiphilic copolymer micelles have a distinct advantage over those formed from conventional surfactants such as Cremophor! EL or polysorbates, since they typically not only display lower CMCs, but also in some cases resist disassembly upon dilution due to physical interactions among chains in the micelle core. Due to their nanoscopic size and the nature by which they are formed, micelles are classified as association or amphiphilic colloids, but should not be considered solid particles (Martin, 1993). It has been shown in several experiments using light scattering, sedimentation velocity and small angle X-ray scattering that the individual molecules or unimers that make up the micelle are in a dynamic equilibrium with the unimers in the bulk and can therefore obey what is termed a closed association model (Alexandridis and Hatton, 1995; Alexandridis et al., 1994; Allen et al., 1999; Tuzar and Kratochvil, 1993).  21  Micelles typically have diameters ranging from 10–50 nm and are characterized by a core-shell architecture in which the inner core is composed of the hydrophobic regions of the amphiphiles, creating a cargo space for the solubilization of lipophilic drugs (Allen et al., 1999; Jones and Leroux, 1999; Kwon, 1998; Kwon and Kataoka, 1995; Lavasanifar et al., 2002). The core region is surrounded by a palisade or corona composed of the hydrophilic blocks of the amphiphiles. The hydrophilic blocks forming the corona region become highly water bound and adopt a “splayed” appearance, giving rise to different conformations such as a polymer “brush” (Kwon and Kataoka, 1995). These conformations sterically suppress opsonization by blood components, thus resisting phagocytosis by macrophages and decreasing clearance by the reticuloendothelial system (RES), resulting in prolonged circulation times (Allen et al., 1999; Kataoka et al., 2001; Kwon, 1998; Kwon and Kataoka, 1995; Lavasanifar et al., 2002).  1.2.3.4 Nanospheres A polymeric nanosphere may be defined as a matrix-type, solid colloidal particle in which drugs are dissolved, entrapped, encapsulated, chemically bound or adsorbed to the constituent polymer matrix (Ameller et al., 2003; Gref et al., 1995; Soppimath et al., 2001). These particles are typically larger than micelles, often having diameters up to 200 nm and may also display considerably more polydispersity (Table 1.2) (Kwon, 1998).  Even though elimination may be slowed by the submicron particle size of nanospheres, clearance is still inevitable due to capture by the RES, sequestering particles within  22 organs such as the liver and spleen (Illum and Davis, 1984). It has been shown that the hydrophobic surfaces of these particles are highly susceptible to opsonization and clearance by the RES. Hence, it became clear that in order to prolong the circulation of nanoparticles, the surfaces must be modified to “look like water” so that they appear to be invisible to the RES (Allen, 1994). Attempts have been made to alter the surface of nanoparticles by adsorbing various surfactants to the particle surface including poloxamine, poloxamer and Brij (Mueller and Wallis, 1993; Troester and Kreuter, 1988; Troester and Kreuter, 1992). Although surfactant coating reduced the total uptake by the RES organs over short periods of time, no difference between uncoated and coated particles was found over longer periods likely due to desorption of the surfactant (Douglas et al., 1987; Gref et al., 1995; Illum and Davis, 1984; Illum et al., 1986). Nanospheres prepared using amphiphilic copolymers such as MePEG-b-PDLLA with high molecular weight hydrophobic blocks provided conjugated PEG coatings with greater stability (Gref et al., 1994; Peracchia et al., 1997). Diblock copolymer nanospheres show a phase-separated structure with a solid core (Gref et al., 1994).  As will be discussed in greater depth below, a clear distinction between micelles and nanospheres formed from diblock copolymers is not always possible, nor desirable. Comprehensive studies using a series of MePEG-b-PDLLA copolymers by Riley et al. and Heald et al. investigated the effects of increasing hydrophobic block length on the physicochemical properties of nanoparticles formed (Heald et al., 2002; Riley et al., 2001). They showed that aggregation behavior and copolymer architecture of MePEG-b- PDLLA was strongly dependent on copolymer composition. As the molecular weight of  23 the PDLLA block increased, the central core of the nanoparticles became more solid-like, resembling nanospheres, whereas smaller PDLLA blocks produced nanoparticles that were termed micelle-like assemblies (Riley et al., 2001).  24 Table 1.2 Nanoparticulate drug delivery systems formed by amphiphilic block copolymers and their general characteristics Nanoparticle Structure Characteristics Micelle  Typically below 50 nm in diameter. Aggregated copolymer chains in dynamic equilibrium with unassociated unimers. Mobile “fluid-like” core. Nanosphere  Typically below 200 nm in diameter. Copolymer chains in a “frozen” state. Phase separated “solid-like” matrix core. Polymersome  Diameter between 5 nm - 5µm Copolymer chains assembled in a bilayer around  an aqueous reservoir. Nanocapsule  Diameter between 100 – 300nm. Copolymer chains or membrane surrounding a drug reservoir or oily core.   25 1.2.4 Rationale for the use of Nanoparticulate Drug Delivery Systems Nanoparticulate drug delivery systems have attracted an extraordinary amount of attention due the advantages they provide for delivering their drug payload. The use of nanoparticles has found particular application in the delivery of anticancer drugs. As will be discussed below, these systems can be utilized to favorably alter the pharmacokinetics and biodistribution of encapsulated drugs via use of passive and active targeting strategies. They have been shown to effectively increase the aqueous solubility of hydrophobic agents. Additionally, there has been great interest in the ability of nanoparticles to overcome multidrug resistance.  1.2.4.1 Prolonged Circulation and the Enhanced Permeation and Retention Effect The tissues of most solid tumors as well as granuloma, inflammatory and infected areas differ from healthy tissues in that they display elevated levels of vascular permeability factors (Maeda et al., 2000). When tumor cells cluster together, angiogenesis is induced to provide the growing multiplying cells with an adequate blood supply. The newly formed vessels differ from those of normal tissues in that they are irregularly shaped, dilated, leaky and the endothelial cells are misaligned and contain large fenestrations (Iyer et al., 2006). Within these vessels there is production of vascular permeability factors such as bradykinin, nitric oxide, vascular permeability factor and prostaglandins (Greish, 2007). Often the smooth muscle layer and basement membrane are missing and the tumor tissues have poor lymphatic drainage. Maeda and coworkers found that high molecular mass proteins such as plasma albumin, transferrin and IgG concentrated in tumor tissues (Matsumura and Maeda, 1986). Interestingly, this does not occur with  26 smaller molecular weight compounds as these substances rapidly diffuse back into the circulating blood and are subject to renal clearance (Seymour et al., 1995). The major reason that larger molecules concentrated in tumor tissue was due to the poor lymphatic drainage of tumors. This finding led to the coining of the term “enhanced permeability and retention” (EPR) to describe the phenomenon (Matsumura and Maeda, 1986). However, not all macromolecular drugs are subject to the EPR effect. It was proposed that there are two prerequisites for the occurrence of the EPR effect. Firstly, the drug must have a sufficient molecular mass to prevent diffusion back into the general circulation. It was found that the threshold molecular weight for polymer conjugated drugs (a.k.a. polymeric drugs) was > 40 kDa. Secondly, the plasma concentration of the drug must be retained at relatively high levels for at least 6 hours. It was envisioned that the EPR effect could be exploited as a passive targeting mechanism for anticancer therapy and since its discovery, several macromolecular anticancer agents have been investigated. One of the first groups of therapeutic agents were polymer drug conjugates, including neocarzinostatin conjugated to copoly(styrene-maleic-acid) also known as SMANCS (Maeda et al., 2001). SMANCS was the first macromolecular anticancer agent and was approved for clinical use in Japan in 1994 for treatment of hepatocellular carcinoma. When administered arterially, SMANCS accumulated in the tumor with extraordinary efficiency due to the EPR effect. Other polymeric drug successes include PEG-conjugated oxidoreductases such as xanthane oxidase and D-amino acid oxidase (Fang et al., 2002; Sawa et al., 2000). Formulations in which the drug is solubilized, encapsulated or bound to a nanoparticle are also subject to the EPR effect. One such well documented success story is the liposomal formulation of doxorubicin, Doxil!.  27  As noted previously, in order for a delivery system to take advantage of the EPR effect, the drug must remain circulating in the blood for a sufficient amount of time. Upon injection of nanoparticulate delivery systems, they are subject to rapid elimination from the blood by the reticuloendothelial system (RES) also referred to as the mononuclear phagocyte system (MPS). The binding of opsonins such as complement protein C3b, immunoglobulin G and M, fibronectin and apolipoproteins to the surface of nanoparticles mediate the recognition and uptake by macrophages (Vonarbourg et al., 2006). The strategy to decrease the rapid elimination of these delivery systems is to prevent the binding of the opsonins to the nanoparticle surface, which has been approached in several ways. The small size of nanoparticles is advantageous and some studies have shown a size dependence in the clearance of nanoparticles, with smaller particles circulating for longer in the blood. It has been suggested that this observation is due to the higher degree of curvature of the surface of smaller particles (Vonarbourg et al., 2006). In order for a complement protein to bind to the surface of a particle it must be in the proper geometric configuration, which is more difficult to achieve on a highly curved surface. Furthermore, since complement activation involves the binding of several components, smaller particles possess less surface area for this binding to occur (Vonarbourg et al., 2006). One of the most effective ways to elude the RES is to coat the surface of the nanoparticles with hydrophilic or charged polymers to prevent the binding of opsonins (Benoit, 2005).   28 1.2.4.2 Platforms for Active Targeting Although passive targeting is effective in some formulations, it also has its limitations. Not all tumors exhibit the EPR effect, and in those that do, the permeability of the vessels may not be the same throughout the tumor (Peer et al., 2007). The specificity of anticancer drugs may be improved with the use of active targeting methods. In this strategy, anticancer drugs are specifically targeted to cancer cells or the tissues associated with tumors, such as the vasculature. Active targeting may be achieved by associating the drug with molecules that bind to antigens or receptors on cells that are uniquely or highly expressed by the target cells.  The use of nanoparticles as a reservoir for drugs with a targeting moiety on the surface has several advantages over conjugating the ligand directly to the drug. By utilizing a nanoparticulate drug carrier, it is possible to significantly increase the ratio of drug to targeting molecules, with the possibility of delivering thousands of drug molecules per targeting molecule. Additionally, nanoparticles can significantly alter the pharmacokinetics of the encapsulated drug compared to the free drug, act as sustained release systems and protect the drug from degradation. Direct conjugation of the targeting moiety to the drug frequently results in decreased activity of the drug (Brannon-Peppas and Blanchette, 2004). Typically, the structure of drugs encapsulated in nanoparticles is unaltered thus preserving the activity of the therapeutic. Several targeting moieties have been investigated including, antibodies or antibody fragments (Park et al., 2001), transferrin (Sahoo and Labhasetwar, 2005), folate (Leamon and Reddy, 2004), lectins and aptamers (Farokhzad et al., 2004), which have been conjugated to the surface of a variety  29 of nanoparticles including liposomes and polymeric nanospheres. Once the targeted nanoparticle binds to its receptor, the drug may elicit its action by one of two mechanisms (Peer et al., 2007). If the receptor is a non-internalizing ligand, the nanoparticle remains on the outside of the cell and releases the drug payload for uptake by free drug transport mechanisms. Alternatively, if the ligand is an internalizing ligand, such as folate, the nanoparticle will be endocytosed and the drug released in the endosome (Parveen and Sahoo, 2008).  1.2.4.3 Drug Solubilization It has been estimated that the number of poorly water soluble potential drug candidates has risen sharply to the order of 40% of all new chemical entities, thus presenting one of the most frequent and greatest challenges for drug development (Neslihan Gursoy and Benita, 2004). The advent of combinatorial chemistry and high throughput screening has allowed for the selection of lead compounds that are excellent ligands, but poor potential drug candidates due to their high molecular weights and increased lipophilicity (Lipinski et al., 1997). The poor water solubility of these compounds can hinder or even prevent the progress of the drug into clinical use. Low molecular weight surfactants such as polysorbate 80 and polyethyloxylated castor oil (Cremophor! EL) are often used to solubilize hydrophobic drugs to enable intravenous injection. Although these excipients are effective at increasing the aqueous solubility of the drug, they have been shown to be biologically and pharmacologically active (Van Zuylen et al., 2001). In the case of Cremophor! EL, significant hypersensitivities and toxicities have been identified (Dorr, 1994). Thus, there has been considerable interest in the use of polymeric nanoparticulate  30 delivery systems composed of biocompatible, biodegradable amphiphilic diblock copolymers, as these systems have been shown to effectively increase the aqueous solubility of several hydrophobic drugs.  1.2.4.4 Evasion of Multi Drug Resistance Mechanisms A significant problem in cancer treatment is the occurrence of multi drug resistance. Cancer cells are capable of developing resistance to chemotherapeutic agents by a variety of mechanisms; however, overexpression of the drug efflux protein, P-glycoprotein (P- gp) is the most frequent (Endicott and Ling, 1989). This membrane bound protein acts a pump that is capable of the efflux transport of numerous drugs, resulting in the reduction of intracellular drug levels and decreased therapeutic efficacy (Gottesman, 2002). Aside from cancer cells, P-gp is also readily found in the apical membrane of enterocytes in the gastrointestinal tract, reducing the oral bioavailability of many drugs (Hunter and Hirst, 1997). Additionally, P-gp is present in the endothelial cells of brain microvesssels and has been implicated in the reduced transport of drugs across the blood brain barrier (Miller et al., 1997).  Several investigators have shown that drug loaded nanoparticles may be used to overcome drug efflux by P-gp (De Verdiere et al., 1997; Kabanov et al., 2002; Nori et al., 2003). Several mechanisms for evasion of P-gp efflux by nanoparticles have been proposed and is an area of intensive investigation. Wong and coworkers demonstrated that drug efflux could be avoided by encapsulating doxorubicin in solid lipid hybrid nanoparticles (Wong et al., 2006). It was found that the nanoparticulate drug was taken  31 up in P-gp overexpressing cells by phagocytosis and retained within the cells at higher levels than unencapsulated drug, thus, showing that endocytosis may be an effective strategy for overcoming drug efflux. Similar results were found for drug encapsulated in immunoliposomes that targeted receptor mediated endocytosis (Huwyler et al., 2002; Kobayashi et al., 2007).  Surfactants such as polysorbates, vitamin E TPGS and Cremophor!EL have been shown to modulate P-gp efflux and thus, the materials that are used to formulate nanoparticulate drug delivery systems may be used to inhibit P-gp. Recently, amphiphilic block copolymers including Pluronics! and MePEG-b-PCL have also been shown to increase the uptake of P-gp substrates. Several mechanisms have been proposed. It was suggested by Kabanov and coworkers that the mechanism by which Pluronics inhibited P-gp was a combination of membrane fluidization and depletion of ATP (Kabanov et al., 2002). Zastre et al. demonstrated that low molecular weight MePEG-b-PCL diblock copolymers were also capable of increasing the Caco-2 cellular accumulation of the P-gp substrates by the inhibition of P-gp. Interestingly, the copolymer did not increase membrane fluidity or inhibit P-gp ATPase activity as shown for Pluronics! (Zastre et al., 2007).  1.3 POLYMERIC DRUG DELIVERY Polymers are used as a delivery platform for a large number of small molecule drugs, peptides and proteins. The drug delivery systems may be fabricated as conjugates, particulates, implants, coatings, films, membranes and other geometries. These drug delivery systems may possess advantages such as: (1) the continuous maintenance of  32 drug levels in a therapeutically desirable range; (2) the reduction of side effects by the targeted or local delivery to a particular cell type or disease site; (3) reduced amount of drug needed; (4) decreased number of doses leading to improved compliance; (5) the facilitation of drug administration of pharmaceuticals with short half-lives (Langer, 1998). These advantages may be achieved by either the controlled release of a drug and/or the targeted or localized delivery of a drug to a particular disease site or tissue.  1.3.1 Polymer Molecular Weight One of the features that distinguishes polymers from small molecules is the lack of a single defined molecular weight. Since polymerization reactions occur through a sequence of random events, the resulting products are characterized by heterogeneous molecular weights and thus, a distribution of molecular weights (Cowie, 1973). Therefore, the molecular weight of a polymer should always be referred to as the average molecular weight. The distribution of molecular weights can be described by several different average molecular weights: the number averaged molecular weight ( ! Mn ), the weight averaged molecular weight ( ! Mw), the z averaged molecular weight ( ! Mz) and the viscosity averaged molecular weight ( ! M"), each of which are obtained by different analytical methods. Since each of these average molecular weights gives a different value it is possible to use several average molecular weights to adequately describe the molecular weight distribution of a sample (figure 1.3). In practice the ! Mn  and ! Mw  are most commonly used for polymer characterization.   33 The number averaged molecular weight uses the numerical fraction of molecules in a class as a weighting factor and may be determined by any method that calculates the number of molecules in each weight class (Cowie, 1973). Since colligative properties such as boiling point elevation, freezing point depression and osmometry are dependent on the number of molecules in solution, these methods may be used to determine the ! Mn . Alternatively, end group analysis may be used to determine the ! Mn . The ! Mn  is a typical numerical average and is calculated by the equation: ! Mn = NiMi" Ni"   (1.1) where Ni is the number of molecules with a molecular weight of Mi.   34        Figure 1.3 Representation of a molecular weight distribution curve with the various average molecular weights indicated.  35 Often a polymer property is more dependent on the size or weight of polymer chains as opposed to the number of polymer chains. Such is the case in light scattering (Odian, 2004). If the mass fraction of molecules in a class is used as the weighting factor, the weight averaged molecular weight ( ! Mw) may be obtained. The ! Mw  may be calculated by the following equation:  ! Mw = wiMi" wi"   (1.2)  where wi is the weight of molecules of molecular weight Mi. Experimentally, the ! Mw  is determined by light scattering. Since the mass in a particular weight class is given by the product of the number of molecules in the class and their molecular weight:  ! wi =NiMi   (1.3)  The ! Mw  can therefore be calculated by:  ! Mw = NiMi 2" NiMi"  (1.4)  Other average molecular weights can be determined by generalizing this formula in which the term Ni is replaced by NiMi k . This gives an average molecular weight ! Mk . When k = 2 the average molecular weight is called ! Mz calculated by the equation:  36  ! Mz = NiMi 3" NiMi"   (1.5)  The ! Mz is determined by ultracentrifugation. The viscosity averaged molecular weight ( ! M") does not follow the formula for ! Mk  but rather is calculated by:  ! M" = (NiMi 1+# ) 1 #$ NiMi$   (1.6)  Where ! is a constant that is dependent on the volume of the solvated polymer in solution. ! M"  is determined by measurements of polymer solution viscosity.  The breadth of the distribution of molecular weights is referred to as the polydispersity index and is calculated as the ratio of ! Mw  to ! Mn  (Rosen, 1982). Since molecules with high molecular weights contribute more to the calculated average when the weighting factor is the mass fraction as compared to the number fraction, ! Mw  is always greater than ! Mn . Therefore, the polydispersity of a polymer sample is always >1. The closer the polydispersity is to unity, the narrower the distribution of the molecular weights of the sample.   37 The molecular weight and molecular weight polydispersity may be determined by gel permeation chromatography (GPC) (Odian, 2004). In this technique, polymer samples dissolved in an appropriate solvent are separated according to their size through a chromatographic column containing a stationary phase composed of highly porous microparticles. Small molecules enter into the pores and spend more time associated with the stationary phase compared to molecules with a higher molecular weight. Detection of polymer mass is accomplished by refractive index and/or light scattering. The instrument is calibrated using a series of low polydispersity polymer standards such as polystyrene or poly(ethylene glycol). The molecular weights ( ! Mnand ! Mw) and the polydispersity of the sample are determined using calibration curves of the log(MW) versus the retention time (Odian, 2004).  1.3.2 Polymer Morphology  It was long believed that polymers could not exist in the crystalline form. This theory was disproved in the 1920s when X-ray diffraction studies demonstrated that elongated cellulose was in fact crystalline (Seymour, 1981). Polymer samples are rarely completely crystalline and are better described as semi-crystalline materials. This phenomenon is displayed in the X-ray diffraction patterns of polymers which are typically characterized by sharply defined rings, indicative of considerable long range order of molecules or sections of molecules, superimposed on a diffuse halo due to the amorphous content (Cowie, 1973). Thus, semi-crystalline polymers are composed of regions of crystallinity, referred to as crystallites, associated with amorphous regions and in fact even highly  38 crystalline polymers are characterized by some amount of amorphous content. Compared to the extended length of a polymer chain, a crystallite is relatively small, leading to the incorporation of a single polymer chain into more than one crystallite. This configuration imparts strain on the crystallite inhibiting its continued growth, leading to small regions of crystallinity. Additionally, due to the high viscosity of a polymer melt, the diffusion coefficient of polymer segments is low, further preventing the formation of large crystallites (Seymour, 1981).  The first model created to describe polymer crystalline structure is termed the fringed micelle model (figure 1.4). In this model, polymer crystallites, or fringed micelles, are composed of regularly arranged portions of polymer chains lined up parallel to one another. Since the chains are many times longer than the crystallites, the chains may pass from one crystallite to another through amorphous regions of randomly arranged polymer chains. This model is useful in describing the coexistence of crystalline and amorphous regions and the increase in crystallinity when polymer fibers are drawn. However, the fringed micelle model has been disproved through the observation of the growth of single polymer crystals (Rosen, 1982).  39                     Figure 1.4 Schematic representation of the fringed micelle model of polymer crystallinity.  40 It is possible to grow single polymer crystals from dilute solution. Through this process, thin, pyramidal or plate-like crystals may be formed termed lamellae. X-ray diffraction of these crystals has shown that the polymer chains are aligned perpendicular to the flat surfaces of the crystals even though the extended length of the polymer chains is greater than the thickness of the crystal (Rosen, 1982). The only way that this could happen is if the polymer chains folded back on themselves to reenter the crystal structure (Rosen, 1982). The resulting model was termed the chain folded model of polymer crystallinity. There is still a considerable amount of amorphous material associated with these crystals, which must be located on the surface due to the disordered arrangement of chains emerging and reentering the crystal structure. Three models of the reentry of chains have been proposed for the chain folded model(Cowie, 1973) (figure 1.5): 1) Regular - reentry folds are adjacent to one another. 2) Irregular - the reentry of chains is adjacent to one another but the folds are of different lengths. 3) Switchboard - reentry of chains is random.  41    Figure 1.5 Models of reentry of chains for the chain folded model of polymer crystallinity. A) regular reentry, B) irregular reentry and C) switchboard reentry.   42  1.3.3 Controlled Drug Release Systems A controlled release system is capable of delivering the drug to the systemic circulation in a predetermined, predictable and reproducible fashion (Sinko and Kohn, 1993). The primary advantage of a controlled release delivery system is that the drug is released such that its blood levels are maintained within the therapeutic window for prolonged periods of time. It is not always beneficial to have prolonged or sustained release of a drug. In the case of diseases such as diabetes, where the drug regimen should ideally mimic the body’s natural cycle, the drug should be delivered in a pulsatile fashion. Therefore, a controlled drug release system may be fully described as one that not only releases drug in a sustained fashion, but may also do so temporally (Sinko and Kohn, 1993). Two fundamental types of controlled drug release systems exist, the matrix system and the reservoir system (figure 1.6). In a matrix system, the drug is uniformly dispersed within a polymer, whereas a reservoir system consists of a separate drug phase surrounded by a rate-limiting polymer phase. Drug release from these systems typically occurs by three mechanisms, solvent controlled release, diffusion controlled release, and chemical controlled release (Sinko and Kohn, 1993).  43     Figure 1.6 Schematic of controlled drug release systems. A) Reservoir system and B) matrix system. Arrows represent transport of drug out of the device.  44  1.3.3.1 Solvent Controlled Release Solvent controlled release may occur by two mechanisms: polymer swelling and osmosis. The most common type of system that undergoes a swelling mechanism of drug release is a hydrogel. These systems are comprised of water-soluble polymers that are rendered insoluble by crosslinking and therefore have the ability to swell in water and retain water within their structure. The rate of swelling, which is determined by the hydrophilic- hydrophobic balance of the polymer and the degree of crosslinking, controls the rate of drug release (Kost, 1995). Osmotic systems are characterized by a drug reservoir surrounded by a non-degradable semi permeable membrane that is permeable to water but not the encapsulated drug. As water enters the system and pressure builds in the device, eventually the drug escapes through a hole in the polymer membrane. An example of this type of device is the osmotic pump tablet (Sinko and Kohn, 1993).  1.3.3.2 Diffusion Controlled Release The rate of drug release from reservoir systems is controlled by the rate of diffusion of the drug through the rate determining membrane. An example of a commercially available reservoir drug delivery system is the transdermal patch such as Transderm- Nitro! (Sinko and Kohn, 1993). In matrix systems, the drug is uniformly distributed within the polymeric phase. If the drug is dissolved in the polymer, the system is referred to as a monolithic solution. Release of drug from a matrix is then described by Fick’s first law:  45 ! J = "D dc dx    (1.7) where J is the flux (given by the rate of diffusion per unit area), D is the diffusion coefficient and dc/dx is the concentration gradient. Equations for the release of the drugs from monolithic solutions with a variety of geometries have been derived (Baker, 1987). Drug release from a slab is described by: ! Mt M0 = 4 Dt "l2 # $ % & ' ( 1/ 2    for 0 ) Mt M0 ) 0.6   (1.8) and ! Mt M0 =1" 8 #2 e "#2Dt l 2 $ % & ' ( )    for 0.4 * Mt M0 *1.0 (1.9) where M0 is the initial amount of drug loaded into the device, Mt is the amount of drug remaining in the device at any given time t, D is the diffusion coefficient and l is the thickness of the device. Release from a device with a spherical geometry may be described by the equations: ! Mt M0 = 6 Dt r 2" # $ % & ' ( 1/ 2 ) 3Dt r 2    for Mt M0 < 0.4   (1.10) and, ! Mt M0 =1" 6 #2 exp "#2Dt r2 $ % & ' ( ) Mt M0 > 0.6   (1.11) where r is the radius of the device. If the drug has limited solubility in the polymer and is dispersed as solid particles, the system is referred to as a monolithic dispersion (Baker, 1987). When the drug loading is low (0-5%) drug release involves the initial dissolution of the drug into the polymer followed by diffusion to the surface of the device. This is  46 referred to as a simple monolithic dispersion and drug release from these devices with a spherical geometry are described by the equation: ! 3 2 1" Mt M0 # $ % & ' ( 2 / 3) * + , - . " Mt M0 = 3DCs(m ) ro 2 C0 t   (1.12) where Cs(m) is the solubility of the drug in the polymer matrix, ro is the radius of the sphere and C0 is the total concentration of drug in the matrix (dissolved and dispersed). For dispersions with a higher drug loading (5-10%) the release mechanism becomes more complex. As material is lost from the surface of the device, cavities are formed which are filled with fluid from the external environment. These fluid filled cavities enhance drug release from the device. This type of device is referred to as a complex monolithic dispersion (Baker, 1987). At drug loadings exceeding 20% the cavities become interconnected to form fluid filled channels through which the majority of the drug is released. These devices are called monolithic matrix systems (Baker, 1987). If the polymer used is non-degradable, the release will be strictly dependent on the diffusion of the drug through the polymer. If degradable polymers are used, the rate of drug release can be dependent on both diffusion of the drug through the matrix and degradation of the polymer.  1.3.3.3 Chemically Controlled Release The rate of drug release from chemically controlled systems is controlled by the rate of polymer degradation. Degradation of the polymer may occur by one of two modes, bulk degradation and surface degradation (figure 1.7). During bulk degradation, the influx of water into the polymer matrix is faster than the breakdown of the material into water-  47 soluble components (Sinko and Kohn, 1993). Therefore, bulk degradation occurs throughout the whole matrix. The result of bulk degradation is the disintegration of the matrix into smaller and smaller pieces. The majority of biodegradable polymers, including the polyesters, degrade by bulk degradation. During surface degradation, the rate of water ingress into the matrix is slower than that of the degradation of the polymer, therefore, the structural integrity of the matrix is maintained, as only the surface undergoes degradation. Very few biodegradable polymers are subject strictly to surface degradation. Two examples of such polymers are polyanhydrides and poly(ortho esters) (Sinko and Kohn, 1993)  48        Figure 1.7 Schematic of bulk and surface degradation mechanisms of polymer degradation during controlled drug release.  49  1.3.4 Targeted Drug Delivery Targeted drug delivery systems are those that accumulate in the tissues of interest. The use of targeted delivery systems allows for the improvement of the therapeutic index by preventing the distribution of the drug to non-target tissues, thus, preventing or reducing the occurrence of toxicities. These systems may be categorized as either passively or actively targeted systems. In passive targeting, the body’s inherent mechanisms are used to direct the delivery system to the site of action. An example of this type of targeting is the passive targeting of microparticulates to the liver by the natural clearance by the reticuloendothelial system (Lecaroz et al., 2007). Actively targeted systems utilize targeting moieties to enable the systems to seek their desired site of action. A popular method of achieving this is to conjugate to the surface of the delivery system biological molecules such as proteins or antibodies which recognize and bind to the targeted site (Dinauer et al., 2005; Zhang et al., 2008).  1.3.5 Polymers for Drug Delivery A biomaterial may be defined as any material that is intended to interface with biological systems to evaluate, treat, augment or replace any tissue organ or function in the body (Williams, 1999). Classes of biomaterials that have been used or are currently in use include metals and metal alloys, ceramics, natural tissues and polymers. Polymers used as biomaterials may be categorized as biodegradable or non-degradable. In the past decades there has been a shift from the use of biostable (ie non-degradable) materials to biodegradable materials. These materials find utility as non-permanent implants as it is  50 not necessary to surgically remove the device at the end of their lifetime. Additionally, the degradation products are often derivatives of endogenous molecules and are therefore non-toxic and well tolerated (Holland and Tighe, 1992). When considering a biodegradable material for medical use, several properties must be taken into account as follows (Nair and Laurencin, 2007): • The material should not cause an inflammatory or toxic response for a prolonged period of time upon implantation in the body • The material should have an acceptable shelf life • The degradation time of the material should match the healing or regeneration process • The material should have appropriate mechanical properties for the intended application and the variation in mechanical properties with degradation should be compatible with the healing or regeneration process. • The degradation products should be non-toxic and readily cleared from the body • The material should have appropriate permeability and processibility for the intended application.  1.3.5.1 Biodegradable Polymers for Drug Delivery In response to the development of new biomedical technologies such as tissue engineering, regenerative medicine, gene therapy, controlled drug delivery and bionanotechnology, there has been an increased demand for the development of new biodegradable polymers (Nair and Laurencin, 2007). Biodegradation of polymers occurs by the cleavage of bonds that are hydrolytically or enzymatically sensitive, leading to  51 erosion of the material (Nair and Laurencin, 2007). Categorization of biodegradable polymers can be based on their source and may be classified as either natural or synthetic. Natural polymers generally degrade by biological processes, whereas synthetic polymers are typically sensitive to hydrolysis. There are several groups of biodegradable synthetic polymers including aliphatic polyesters, poly(ortho esters), poly(!-amino acids) and polyanhydrides. Natural polymers include polysaccharides and proteins. The most widely used synthetic polymer biomaterials are the aliphatic polyesters including poly(glycolic acid), poly(lactic acid), poly(lactic-co-glycolic acid) and poly(caprolactone) (Piskin, 1995). These polymers are well known for their biocompatibility and controllable degradation profiles. Degradation of these materials is mainly by bulk erosion.  Biodegradation of polymers via hydrolysis is affected by several factors. The hydrophobicity of the polymer plays a role in the degradation rate as is evidenced by the slower degradation rate of poly(caprolactone) as compared to poly(glycolic acid). Polymer morphology influences degradation since water cannot penetrate into the crystalline regions of semi crystalline polymers as readily as amorphous regions, therefore amorphous regions are degraded faster. Amorphous components of the polymer degrade faster if the glass transition temperature is below body temperature. At the glass transition temperature, the amorphous regions of the polymer chains transition from a glassy state to a rubbery state due to increased motion of the polymer chains (Odian, 2004). Above this temperature the rate of diffusion of water into the bulk polymer increases leading to faster degradation (Amsden, 2007). The geometry of the system also influences the degradation rate, with systems characterized by high surface area to  52 volume ratios degrading more rapidly due to an increase in the rate of uptake of water. In the case of polyesters, hydrolytic cleavage of ester bonds produces carboxylic end groups, which are capable of catalyzing the hydrolysis of other ester groups (Li, 1999). The mechanism, termed autocatalysis, can increase the degradation rate of polyesters. Additionally the site where the polymer is implanted or injected will affect the degradation rate as there may be variations in the pH, enzyme activity and temperature (Piskin, 1995).  1.4 PACLITAXEL 1.4.1 Chemistry Paclitaxel (PTX) is a diterpenoid with the chemical formula C47H51NO14 and a molecular weight of 853.9 g/mol. The IUPAC chemical name for paclitaxel is 5!, 20-epoxi- 1,2",4,7!, 10!,13"-hexahydroxytax-11-en-9-one-4,10-diacetate-2-benzoate 13 ester with (2R,3S)-N-benzoyl-3-phenylisoserine. The structure of PTX is composed of a complex taxane ring system bound to an oxetane ring at C-4 and C-5 with an ester side chain at C- 13 (figure 1.8). The compound was first isolated from the bark of the Western Yew, Taxus brevifolia (Wani et al., 1971). Since the extraction process yields very little product with only 0.04% recovered from the dry weight of the bark and the procurement of product would lead to the destruction of the already limited supplies of Taxus brevifolia, alternative methods of obtaining PTX were sought. A semisynthetic approach using a precursor, 10-deacetyl baccatin III, which is readily available from the needles of a related species, Taxus baccata L, was developed by Denis et al. (Denis et al., 1988). PTX is a white crystalline powder that is soluble in organic solvents including methanol,  53 ethanol, acetonitrile and methylene chloride. The aqueous solubility of PTX low and has been reported to range from 0.7 – 30 µg/ml with a log P of 3.5 (Mathew et al., 1992; Swindell et al., 1991). Liggins et al. showed the existence of a dihydrate with an aqueous solubility of 1 µg/ml (Liggins et al., 1997). In aqueous solution, the structure of PTX undergoes epimerization of the C7 hydroxyl group via a base catalyzed mechanism to form the more thermodynamically stable 7-epi PTX (Tian and Stella, 2007). 7-epi PTX has similar activity as the parent compound (Ringel and Horwitz, 1987).  54               Figure 1.8 The chemical structure of paclitaxel. O O O O OH OO O HO NH OH OO O O 13 4 7 10 1 2 3 5 6 8 911 12 14 1' 2' 3'  55  1.4.2 Pharmacology PTX inhibits cell proliferation by different mechanisms. The classic mechanism of action is through the binding to !-tubulin causing the formation of unusually stable microtubules arresting the cell cycle at the G2-M phase (Rowinsky and Donehower, 1995). This mechanism is unique compared to other anticancer drugs that act on mitotic spindles such as vincristine, vinblastine, colchicines and podophyllotoxin, which act to induce the disassembly of microtubules (Schiff et al., 1979). Conversely, PTX inhibits the disassembly of microtubules leading to apoptosis of the cell (Rowinsky et al., 1990). PTX shows activity against a range of cancer cell types including ovarian, breast, lung, prostate, melanoma and leukemia (Rowinsky and Donehower, 1995). It was first approved in 1992 as a first line therapy for the treatment of epithelial ovarian cancer. Since that time it has been approved for use in breast cancer, non small lung cancer and AIDS related Kaposi’s sarcoma (Sparreboom et al., 2005).  1.4.3 Toxicity The main dose limiting toxicity of PTX is neutropenia, which may be severe at doses of 200 to 250 mg/m 2 . It was found that PTX does not irreversibly damage hematopoietic stem cells and duration is brief with complete recovery by day 21 (Rowinsky et al., 1992). One of the most life threatening toxicities associated with PTX administration is hypersensitivity reactions. The most severe of which are type I reactions, which include dyspnea with bronchospasm, urticaria, abdominal and extremity pain, angioedema, and  56 diaphoresis (Rowinsky et al., 1992). It is believed that the excipient used to solubilize the drug, Cremophor! EL, is the culprit for the hypersensitivity reactions as it has been associated with hypersensitivity reactions in other formulations in which it is used as a solubilizer (Dorr, 1994). In order to avoid hypersensitivity reactions, patients are premedicated with steroids, and H1 and H2 antihistamines (Rowinsky et al., 1992). Other toxicities of PTX include peripheral neurotoxicities and cardiac rhythm disturbances (Rowinsky et al., 1992).  1.4.4 Pharmacokinetics PTX is clinically administered as a single agent in 1, 6 or 24 hour infusions with dosages ranging from 15-275 mg/m 2  (Kuhn, 1994). The maximum concentration in plasma (Cmax) is dependent on the length of infusion and the dose, with shorter infusion times resulting in higher Cmax values than for longer infusions (Straubinger, 1995). Upon administration, the drug rapidly distributes to the tissues with a distribution half life (t1/2") of 15 to 30 minutes. The elimination half life (t1/2#) is between 1.3 to 8.6 hours with a mean value of 4.9 hours. Elimination of PTX is primarily by hepatic metabolism, with less than 10% of the drug excreted in the urine and the majority excreted in the bile and eliminated in the feces (Kuhn, 1994; Straubinger, 1995). The steady state volume of distribution ranges from 49 to 119 L/m 2  with an average of 110 L/m 2  indicating binding to plasma proteins and tissues (Straubinger, 1995).   57 1.4.5 Paclitaxel Loaded Nanoparticulate Formulations Due to its extremely hydrophobic nature, the commercial formulation of paclitaxel consists of the drug at a concentration of 6 mg/ml solubilized in polyethoxylated castor oil (Cremophor! EL) with 50% anhydrous alcohol. The trade name for this product is Taxol". Prior to administration, the formulation is diluted 5 – 20 fold with normal saline or 5% dextrose (Straubinger, 1995). Cremophor! EL is efficient at solubilizing paclitaxel and has been reported to remain stable upon dilution for greater than 24 hours (Straubinger, 1995). Cremophor! EL is currently used to solubilize several other intravenously (i.v.) administered drugs such as cyclosporine A and teniposide, however, the concentration needed to solubilize PTX is considerably higher than for these agents. As mentioned previously, administration of Cremophor! EL is associated with several adverse effects including anaphylactic reactions, hyperlipidemia, and modification of electrophoretic and density gradient behavior of lipoproteins (Sparreboom et al., 1998). Furthermore, this excipient is known to be incompatible with poly(vinyl chloride) infusion sets and causes the leaching of plasticizers such as di(2-ethylhexyl) phthalate (Kim et al., 2005). Accordingly, Cremophor! – free formulations for the intravenous delivery of PTX have been investigated including liposomes (Yang et al., 2007) and cyclodextrins (Bilensoy et al., 2008). There has been considerable interest in the use of amphiphilic block copolymers nanoparticles due to their high degree of biocompatibility, ability to solubilize high concentration of the drug and potential to passively target tumor sites by the EPR effect.   58 Zhang et al. developed the first amphiphilic copolymer nanoparticulate drug delivery system for PTX (Zhang et al., 1996). This micellar system was based on the diblock copolymer MePEG-b-PDLLA with a 60:40 weight ratio of MePEG (MW 2000 g/mol) to PDLLA and was demonstrated to dramatically increase the aqueous solubility of the drug with a drug loading of up to a maximum of 25% w/w (Zhang et al., 1996). This formulation was screened for efficacy against a number of cancer cell lines including Hs578T breast, SKEMES non-small cell lung, and HT-29 colon. In vitro cyctotoxicity results indicated that the polymeric micellar formulation was equally as efficacious as Taxol! (Zhang et al., 1997). Toxicity studies indicated that the polymeric micelle formulation was considerably better tolerated, with maximum tolerated doses (MTD) of 100 mg/kg/day and 25 mg/kg/day for intraperitoneal (i.p.) and intravenous administration, respectively, compared to 20 mg/kg/day for Taxol! administered by either route (Zhang et al., 1997). This group assessed the in vivo efficacy and biodistribution of the formulation in an intraperitoneal P388 leukaemia murine tumor model treated by i.p. injection (Zhang et al., 1997) and a subcutaneously implanted MV- 522 lung tumor treated by both i.p and i.v. administration (Zhang et al., 1997). In both cases, i.p. administration of the polymeric micelle formulation was more efficacious than Taxol!due to the large increase in the MTD and thus total dose of paclitaxel given to the animals. The i.v. administration of the formulation was as efficacious as Taxol!. There was no major difference in the biodistribution of the drug when given as either Taxol!or the polymeric micelle form. Ramaswamy et al. compared the in vitro distribution of PTX solubilized in MePEG-b-PDLLA micelles and free PTX into the component fractions of plasma. Both free and micellized PTX rapidly distributed into lipoprotein and lipoprotein  59 deficient plasma fractions suggesting that the drug was released from the micelles and associated with the plasma proteins, although, this could not be confirmed since only the distribution of the drug was followed and it could not be confirmed whether the drug was still associated with the copolymer (Ramaswamy et al., 1997). Further in vivo biodistibution studies were reported by Burt and coworkers in which the biodistribution of both the copolymer and drug was determined after i.v. injection of radiolabelled micelles and drug in male Sprague-Dawley rats (Burt et al., 1999). It was confirmed that the drug rapidly dissociated from the copolymer and distributed widely to the majority of tissues, whereas the copolymer was rapidly eliminated in the urine.  Since the development of this formulation, many other paclitaxel copolymer nanoparticulate systems have been investigated, utilizing a wide range of materials. Typically, the formulations increased the aqueous solubility of paclitaxel and generally possessed equivalent or better cytotoxicity against tumor cells as compared to Taxol!; however, very few studies have investigated the in vivo biodistribution and pharmacokinetics of these novel formulations. Le Garrec and co workers evaluated the in vivo performance of paclitaxel loaded PVP-b-PDLLA micelles in a murine colon adenocarcinoma C26 model (Le Garrec et al., 2004). Similar to the findings of Zhang et al., the drug was rapidly released from the micellar carrier resulting in a shorter elimination half life and lower area under the curve compared to Taxol! and the only advantage to this system was the marked increase in the MTD compared to Taxol!. Kim et al. also evaluated a PTX loaded MePEG-b-PDLLA micellar formulation and showed an increase in the MTD, allowing for a greater dose of the drug to be administered but  60 similar to Le Garrec et al., this formulation also had a shorter elimination half life and lower area under the curve than Taxol! (Kim et al., 2001).  In order to increase the retention of PTX within the micellar core, several strategies have been developed. Kim et al. synthesized MePEG-b-PDLLA copolymers with polymerizable methacryoyl end groups and subsequent end-capping with methacrylic anhydride. After micellization the core blocks were polymerized chemically and photochemically. It was demonstrated that the micelles could increase the aqueous solubility of PTX and were stabilized against dissociation in vitro (Kim et al., 1999). Similarly, core cross-linked MePEG-b-PCL diblock and triblock copolymer micelles effectively solubilized PTX and were stable against dilution in water (Shuai et al., 2004). Hydrotropic polymeric micelles have shown promise by solubilizing large quantities of PTX and displaying prolonged stability in aqueous solution but have not been evaluated in vivo (Huh et al., 2008; Lee et al., 2007). Jing and coworkers conjugated PTX via the 2’ or 7- hydroxyl groups of the drug to di or triblock MePEG-b-PLLA copolymers as well as a novel copolymer, poly(lactic acid)-co-[(glycolic acid)-alt-(L-glutamic acid)]-block- poly(ethylene glycol)-block-poly (lactic acid)-co-[(glycolic acid)-alt-(L-glutamic acid)] (Xie et al., 2007; Xie et al., 2007; Zhang et al., 2005). The drug was released from the conjugate and showed in vitro cytotoxicity against cancer cell lines. By increasing the interaction between the hydrophobic block and the solubilized drug, it has been shown that it is possible to retain PTX within the micelle for extended periods of time. Hamaguchi et al. used a PEG-b-polyaspartate diblock copolymer which they modified by bonding 4-phenyl-1-butanolate to the carboxylic acid groups of the polyaspartate block.  61 This allowed for an increased hydrophobic interaction between PTX and the copolymer. This formulation displayed longer circulation in the blood, and preferential uptake into HT-29 human colorectal tumors in a mouse xenograft model. The formulation was reported to be less neurotoxic that free PTX (Hamaguchi et al., 2005). This formulation is currently in clinical trials under the trade name NK105. Using a different approach, Forrest et al. altered the structure of PTX to form prodrugs that were more compatible with the core of MePEG-b-PCL micelles and displayed a greater micellar drug loading (Forrest et al., 2008). A lead candidate with a high drug loading and cytotoxicity equivalent to PTX was chosen for pharmacokinetic and biodistribution studies. The formulation displayed a longer plasma half life and area under the curve and reduced the volume of distribution compared to free prodrug and Taxol!. The in vivo efficacy and toxicity are still to be investigated.  1.5 PHYSICOCHEMICAL PROPERTIES, DRUG SOLUBILIZATION, AND HEMOCOMPATIBILITY OF AMPHIPHILIC BLOCK COPOLYMER NANOPARTICULATE DRUG DELIVERY SYSTEMS 1.5.1 Methods of Polymeric Nanoparticle Preparation and Drug Loading The method used to form drug loaded polymeric nanoparticles is mainly dependent on the aqueous solubility of the polymer used. There are four methods of preparing nanoparticles: direct dissolution, film casting or solvent evaporation, dialysis or nanoprecipitation and emulsification (figure 1.9).   62     Figure 1.9 Methods of micelle preparation and drug loading. A) direct dissolution, B) nanopreciptation or dialysis, C) film casting or solvent evaporation D) emulsification  63 1.5.1.1 Direct Dissolution If the copolymer to be used is readily water soluble, micelles may be formed by simply dissolving the copolymer above the CMC in an aqueous phase. As long as the copolymer concentration remains above the CMC, micelles will be formed. In order to incorporate hydrophobic drugs into these micelles, the simplest method is to add the solution of preformed micelles to solid drug. The low aqueous solubility of the drug allows for the partitioning of the drug into the micelle core. This method of loading has been used extensively for the loading of the fluorescence probe pyrene for the investigation of the CMC and partition coefficient of various copolymers (Lee et al., 1999; Lele and Leroux, 2002).  1.5.1.2 Film casting or solvent evaporation An alternative method of micellar solubilization of drug using water-soluble copolymers is the film casting method, also referred to as the solvent evaporation technique. In this method, the copolymer and drug are dissolved in an organic solvent. This solution is then aliquoted into a vial and the solvent removed by evaporation. The resulting matrix is copolymer with drug dissolved in it. To form micelles, warm water or buffer is added and the matrix agitated until it dissolves. Several drugs have been encapsulated by this method including paclitaxel (Zhang et al., 1996).   64 1.5.1.3 Dialysis and Nanoprecipitation In cases where the copolymer of interest is not water soluble, a dialysis technique may be employed. Poor aqueous solubility is not a requirement for use of this method as micelles composed of water-soluble copolymers may also be formed by this method, in which the copolymer and drug are dissolved in a water miscible organic solvent. The solvent is then added drop-wise into the aqueous phase, which is stirred rapidly. In order to remove the solvent from the aqueous phase, the dispersion is dialysed against water or buffer. During dialysis it is essential that the molecular cut-off of the membrane is below the molecular weight of the copolymer. In a variation to this method, water is added drop-wise to the organic solvent until particles begin to assemble, as determined by changes in optical density or an increase in light scattering. Again, the solvent is removed by dialysis. If the copolymer that is being used is extremely hydrophobic, the combining of the aqueous and organic phases will induce precipitation of the polymer and a fine dispersion of nanoparticles will result. This method is referred to as nanoprecipitation and is commonly used to form nanospheres (Galindo-Rodriguez et al., 2004; Reis et al., 2006).  1.5.1.4 Emulsification Emulsification techniques have been used to form nanospheres composed of homopolymers with a PEG coating or diblock copolymers with long hydrophobic blocks (Gref et al., 2000). This method involves the dissolution of the polymer and drug in a water immiscible organic solvent with a relatively high vapor pressure. This solution is added drop wise to the aqueous phase that is either being stirred rapidly or sonicated. The aqueous phase may contain a surfactant such as poly(vinyl alcohol), particularly if the  65 polymer being used is a homopolymer. After addition of the organic phase, the solvent is allowed to evaporate, resulting in a dispersion of nanoparticles.  1.5.2 Factors Influencing Physicochemical Properties and Drug Loading 1.5.2.1 Methods of Preparation To date, most studies of the formation of copolymer nanoparticles have focused on how variation in preparation techniques have an influence on the loading of drugs, rather than on the physicochemical properties of the nanoparticles (Kim et al., 1998; La et al., 1996; Yu et al., 2002). In studies conducted by Vangeyte et al., a systematic variation of the preparation technique and solvents used and their influence on the size of resulting MePEG-b-PCL particles was conducted (Vangeyte et al., 2004). It was found that direct dialysis of copolymer and solvent solutions led to the formation of large aggregates indicating a fast exchange of solvent, most likely due to the large porosity of the dialysis membrane used. Nanoparticulate formation by nanoprecipitation was also explored by varying the solvent used and the order of addition (i.e. organic phase added to aqueous phase or vice versa). There was very little difference found between nanoparticles formed from the various solvents, with the exception of DMSO or THF, where the particles were consistently larger. The lack of size difference was suggested to be due to the rapid precipitation of the polymer, locking the polymer chains in a kinetically stable conformation and thus, there was no clear dependence on the compatibility between the constitutive blocks and the solvents as calculated by Hildebrand solubility parameters. The increase in particle size when DMSO or THF were used was thought to be due to slower mixing rates of these solvents with water, caused by higher viscosity and a lower  66 miscibility with water, respectively. When the nanoparticles were formed by adding the organic phase of DMSO or THF to the aqueous phase, larger particles resulted, compared to when the addition order was reversed. This was believed to be caused by the faster rate of precipitation when the solvent was added to the aqueous phase (Vangeyte et al., 2004). Interestingly, in the preparation of nanoparticles of methacrylic acid copolymer (Eudragit L 100-55) with poly(vinyl alcohol) as a surfactant by nanoprecipitation, the mean particle size was clearly dependent on the compatibility of the solvent with water. However, this trend was not seen with the PEG-PCL nanoparticles prepared by Vangeyte et al. (Galindo-Rodriguez et al., 2004).  Studies of the effect of other formulation parameters on the preparation of amphiphilic copolymer nanospheres and nanocapsules have shown that increases in the surfactant concentration, either the copolymer itself or an additional surfactant, resulted in a decrease in the diameter of the particles (de Faria et al., 2005; Gref et al., 1995). If additional surfactants were used, for example poly(vinyl alcohol) or cholic acid, it was shown that they may remain associated with the surface of the particle even after extensive washing, resulting in an alteration in the surface charge of the nanoparticle.  As noted previously, the rapid addition of organic phase to water, or vice versa, results in the almost instantaneous precipitation of the polymer and the kinetic locking of copolymer chains into the formed structure. However, Eisenberg and coworkers have shown that the dissolution of asymmetric crew cut aggregate-forming copolymers in organic solvents with small amounts of water (5.5 - 9.5 wt %) induced the aggregation of  67 the copolymers. The dynamic equilibrium between unimers and aggregates is preserved provided the water content is not too high (Zhang and Eisenberg, 1999). This chain mobility allows for the formation of thermodynamically stable structures with varying morphologies, which can then be “frozen” upon the addition of more water and subsequent dialysis of the remaining organic solvent. The various morphologies formed depend on the water content used for the initial dissolution of the polymer, the copolymer concentration in the organic solvent prior to addition of water and the presence of ions (Zhang et al., 1996). The use of various solvents for the dissolution of the copolymer also affected the morphology of the resulting nanoparticles (Yu et al., 1998). It was shown that the water concentration at which copolymers start to aggregate termed the critical water content (CWC), was dependent on the nature of the solvent. The CWC increased as the compatibility between the solvent and core-forming block increased. Nanoparticle morphologies were found to be a result of a balance between interactions between the solvent and both the core and corona forming blocks, resulting in changes in aggregation number and degree of stretching of the core forming block determining which morphology was most entropically favorable.  1.5.2.2 Block Length In the literature it has been shown that as the ratio of the molecular weights of the hydrophilic and hydrophobic blocks changes, the method of preparation to obtain particular nanoparticle formulations needs to be correspondingly altered (Soppimath et al., 2001). When the molecular weight of the hydrophilic block exceeds that of the hydrophobic block, the copolymer is easily dispersed in water and will self-assemble into  68 small, relatively monodisperse micelles. However, when the molecular weight of the hydrophobic block approaches or exceeds the molecular weight of the hydrophilic block, the copolymer becomes progressively more water insoluble and therefore will not self- assemble into a nanoparticle through direct dissolution or film casting methods, but rather dialysis, emulsification or in some cases nanoprecipitation techniques must be employed.  The fundamental studies of Riley et al. and Heald et al. using MePEG-b-PDLLA copolymers with a range of PDLLA molecular weights and a fixed MePEG molecular weight of 5000 g/mol showed that if the PDLLA molecular weight was relatively low (2000 – 30,000 g/mol), the hydrodynamic radius (Rhyd) of the resulting particles was independent of the concentration of the polymer used during preparation and the polydispersity index was low, characteristic of block copolymer micelles (Heald et al., 2002; Riley et al., 2001). It was determined that the Rhyd and aggregation number (Nagg) of nanoparticles are highly dependent on the length of the constituent blocks. Power laws have been developed to express the dependence of the Rhyd and Nagg of nanoparticles on the hydrophobic and hydrophilic block lengths, designated NA and NB respectively (Foerster et al., 1996; Gao et al., 1994; Halperin, 1987). In the case of star micelles in which NB >> NA, the scaling relationships were found to be: Rhyd ! NA 4/25  NB 3/5  and Nagg ! NA 4/5  or for strongly segregated blocks Nagg ! NA 2   69  For the other extreme, crew cut micelle-like aggregates in which   NB << NA, the scaling relationships were found to be: Rhyd ! NA 2/3  and Nagg ~ NA Riley et al. found that the hydrodynamic radius of particles formed by copolymers with PDLLA molecular weights between 2,000 – 30,000 g/mol scaled linearly when plotted against NA 1/3 . This suggested a micellar regime somewhere between star micelles and crew cut micelle-like assemblies, and they settled on the term “micelle-like” assemblies. In contrast, nanoparticles formed from copolymers with higher PDLLA molecular weight segments (45,000 – 110,000 g/mol) displayed hydrodynamic radii that were dependent on the concentration of copolymer solutions used during preparation. The nanoparticles produced were more “particulate-like” with a “solid-like core” (Heald et al., 2002; Riley et al., 2001), analogous to nanospheres shown in Table 1.2. The aggregation number of particles formed from diblocks with relatively low PDLLA molecular weights increased sharply, scaling in good agreement with micellar assemblies of strongly segregated blocks. Further increases in the PDLLA molecular weight from 9000 - 30,000 g/mol resulted in deviations from the scaling dependence of aggregation number on the PDLLA block length indicative of kinetically “frozen” micellar systems. Additional increases in PDLLA block length resulted in an aggregation number dependence on the concentration of copolymer used in preparation.   70 Heald et al. used a variety of NMR techniques including 1 H liquid state, T1 relaxation and 13 C solid state, to study the shift in the state of the copolymers comprising nanoparticles as the molecular weight of the PDLLA block was increased from 2,000 to 25,000 g/mol at constant MePEG molecular weight (Heald et al., 2002). Solid state NMR experiments confirmed the presence of two phases in the PDLLA core of the nanoparticles formed by all the copolymers regardless of the PDLLA molecular weight: a solid-like core and a more mobile interfacial region. Upon heating the nanoparticulate samples, it was found that the amount of methyl and methine protons detectable in the low molecular weight PDLLA samples (2,000 – 4,000 g/mol) increased indicating a more fluid core, whereas copolymers with PDLLA molecular weights above 6,000 g/mol displayed little change indicating a solid core.  The morphology of prepared amphiphilic block copolymer nanoparticles is typically spherical, particularly if the molecular weight of the hydrophilic block exceeds that of the hydrophobic block, thus forming aggregates in which the corona is larger than the core (so-called star micelles) (Allen et al., 1999). However, if the copolymer is asymmetric in its relative block lengths (i.e. the hydrophobic block is considerably longer than the hydrophilic block) and the nanoparticles are carefully prepared by the slow addition of water to the polymer dissolved in a water miscible organic solvent, varying morphologies can be obtained (Cameron et al., 1999; Heald et al., 2002; Zhang and Eisenberg, 1995). To date, the majority of the work in this field has been done using diblock copolymers composed of hydrophobic polystyrene (PS) and hydrophilic poly(acrylic acid) (PAA) blocks (Yu et al., 1996; Zhang and Eisenberg, 1998; Zhang and Eisenberg, 1995).  71  The self-assembly of amphiphilic copolymers into polymersomes has been found to be dependent on the weight fraction of the hydrophilic block. Discher et al. have stated that in the case of PEG-PBD or PEG-PEE copolymers, if the fraction of PEG (fEO) is between approximately 20% - 42% the copolymers will self assemble into fluid-like, bilayer- forming vesicles. If the copolymer is considerably more hydrophobic with a fEO < 20% the immobile hydrophobic blocks will be sequestered into solid-like particles (referred to in this work as nanospheres). For fEO > 42% typically spherical micelles are formed (Discher and Eisenberg, 2002). Discher and Eisenberg and their groups have established that aggregate morphology is principally determined by time-average molecular geometry of diblock copolymers, such that the balance of hydrophilic/hydrophobic blocks produces molecular shapes of cylinder, wedge or cone and this in turn, dictates whether membrane, rod-like or spherical morphologies will form (Ahmed and Discher, 2004; Discher and Eisenberg, 2002).  It is evident that caution must be used when using terminology for amphiphilic copolymer nanoparticles, as it is apparent that not all core shell nanoparticles are appropriately defined as a micelle. There is some evidence that amphiphilic copolymer nanoparticles are occasionally termed micelles, even though the hydrophobic block length is exceptionally long and clearly not water soluble (Aliabadi et al., 2005; Kim et al., 1998; Shuai et al., 2004). Typically these nanoparticles are formed by nanoprecipitation methods, which do not require self-assembly but rather the “freezing”  72 of the copolymer chains into a kinetically stable structure upon removal of the solvent. The resulting structure is that of a core-shell nanoparticle with a solid-like core.  1.5.2.3 Drug and Polymer Compatibility Many factors may affect the amount of drug solubilized by polymeric nanoparticles, including the hydrophobic block length, the molecular volume of the drug, interfacial surface tension, polarity and hydrophobicity (Allen et al., 1999). The compatibility between the core forming block and the drug is a major determinant influencing the amount of drug solubilized by a polymer (Allen et al., 1999; Lavasanifar et al., 2002; Liu et al., 2004).  In order for a solute to be solubilized by a solvent, the intermolecular forces between the solvent and solute must be great enough to overcome the solvent-solvent and solute- solute interactions. The cohesive energy (Ecoh) is the increase in the internal energy per mole if all its intermolecular forces are eliminated (Bicerano, 2002). The amount of energy required to do so in a unit volume of material (V) is defined as the cohesive energy density (ecoh). Therefore:  ! ecoh " Ecoh V   (1.13)  The Ecoh of low molecular weight liquids is found by determining the heat of evaporation; however, polymers cannot be evaporated therefore, indirect methods must be used  73 (Bicerano, 2002). Experiments such as comparative swelling or dissolution in liquids of known cohesive energy densities may yield Ecoh values for some polymers. Dunkel showed that for low molecular weight substances, the Ecoh is an additive property and it could be predicted by summing the contributions of each group making up the substance (Dunkel, 1928). Hoftyzer and Van Krevelen later applied this knowledge to polymers creating tables of the contributions of functional groups for the prediction of Ecoh of polymers (van Krevelen, 1997).  There are three intermolecular forces that contribute to the cohesion of materials, dispersion forces, polar cohesive forces and hydrogen bonds (van Krevelen, 1997). Dispersion forces are induced when the oscillation of charge in a molecule results in a temporary dipole moment. This dipole interacts with the temporary dipoles in neighboring molecules resulting in cohesion between the two molecules. Polar cohesive forces are due to the interaction of permanent dipoles in neighboring molecules. Hydrogen bonding occurs between the partial negative charge on an electronegative atom on a molecule and a partial positive charge in a hydrogen atom of an adjacent molecule. The total cohesive energy is therefore a sum of the energy from dispersion forces, polar cohesive forces and hydrogen bonding abbreviated as Ed, Ep, and Eh, respectively (van Krevelen, 1997). The total Ecoh is calculated as:  ! Ecoh = Ed + Ep + Eh    (1.14)  74 The Hildebrand solubility parameter (!) is a quantitative value of the strength of the intermolecular forces between solvent molecules or solute molecules (Bicerano, 2002). The ! is related to the ecoh and Ecoh by the relationship:  ! " # ecoh = Ecoh V    (1.15)  The total Hildebrand solubility parameter (!t) takes into account the contributions from dispersion (!d), polar (!p) and hydrogen bonding (!h) components and is calculated by (van Krevelen, 1997):  ! "t 2 = "d 2 + "p 2 + "h 2    (1.16)  The solubility parameter can be used to estimate whether two components are miscible by calculation of the Flory-Huggins interaction parameter ("sp) as calculated by:  ! "sp = Vs(#solute$#polymer)2 RT   (1.17)  where Vs is the molar volume of the solute, R is the gas constant and T is the absolute temperature. According to Flory-Huggins solution theory, the "sp is proportional to the enthalpy of mixing (Bicerano, 2002). The lower the "sp value the more enthalpically favorable the mixing of two components and the more likely that the solute will dissolve in the solvent. A large positive "sp indicates a tendency for phase separation to occur as  75 solute and solvent molecules would prefer to be surrounded by molecules of their own kind (Bicerano, 2002).  1.5.3 Hemocompatibility of Polymer-Based Drug Delivery Systems In order for a biomaterial to be considered biocompatible, the material and its degradation products should not induce undesirable host reactions including thrombosis, inflammation, necrosis, toxicity, allergic reactions, or carcinogenesis (Piskin, 1995). For delivery systems that are intended for intravenous injection, it is important that the systems do not elicit any unwanted effects when in contact with blood components.  1.5.3.1 Coagulation Hemostasis or coagulation is the body’s normal response to preventing the leaking of internal fluids. However, when inappropriate or pathological coagulation occurs it is referred to as thrombosis and may result in the formation of an embolus (Piskin, 1995). The end result of the coagulation process is that soluble fibrinogen is transformed into an insoluble fibrin clot through the influence of the enzyme thrombin. The coagulation process involves a cascade of several enzymes, cofactors and phospholipidic surfaces termed coagulation factors (Hanson, 2004). The initiation of coagulation is divided into two pathways, the intrinsic and extrinsic pathways, both of which terminate in the formation of thrombin (Hanson, 2004). The intrinsic pathway is initiated when the blood comes in contact with an anionic surface and is dependent only on factors that are intrinsic to flowing blood. The extrinsic pathway is initiated by tissue factor, also referred  76 to as TF, thromboplastin or factor III, which is exposed to blood when the vascular wall is damaged (Gorbet and Sefton, 2004). When a foreign surface such as a polymer comes in contact with blood, there is a competitive adsorption of proteins and glycoproteins on the surface (Piskin, 1995). The composition of the adsorbed proteins, along with their conformational changes, plays an important role in the biological behavior of the material (Chauvel et al., 2001).  The induction of coagulation in the presence of an artificial surface such as a polymer is induced by activation of the contact phase. During this process, thrombin is produced by a cascade that is triggered by the contact of an electronegative surface and factor XII (Hanson, 2004). Activation of the enzymatic cascade involves the binding of high molecular weight kininogen, prekallikrein / kallikrein, factor XII and factor XI to the surface of the material. The result is increased amounts of factor XIa, which in turn activates factor IX to factor IXa, which then activates factor X to factor Xa in the presence of phospholipids from platelet membranes and factor VIII and calcium ions (Hanson, 2004). Factor Xa, factor V and calcium ions bind to platelet phospholipids in a complex called prothrombinase to convert factor II (prothrombin) into thrombin, which in turn, converts fibrinogen into fibrin (Hanson, 2004).  1.5.3.2 Complement System Activation The complement system is one of the main defense systems the body has to recognize and clear foreign objects. The complement systems is made up of multiple proteins that have enzymatic functions or act as receptors for cells of the immune system. There are  77 three main initiation pathways of the complement system, the classical pathway, the alternate pathway and the lectin pathway (Vonarbourg et al., 2006). Activation of the classic pathway takes place when complement protein C1 binds to an antigen-antibody complex. This starts a cascade of binding and cleavage events that culminates in the formation of classical C3 convertase (Vonarbourg et al., 2006). The lectin pathway is similar to the classical pathway, with the exception that it is initiated by the binding of mannose binding lectin to mannose residues on the pathogen surface. This activates mannan binding lectin associated serine proteases, which then tie into the classical pathway resulting in the production of classical C3 convertase (Passirani and Benoit, 2005). The alternate pathway is a humoral component which operates without antibodies. This pathway is activated by the hydrolysis of C3 into C3a and C3b. The latter binds to the surface of the pathogen forming alternative C3 convertase. C3 convertase produced by either the classical or the alternative pathways is cleaved into C3 which then starts a cascade that results in the formation of the membrane attack complex (MAC). The MAC forms pores in the surface of the pathogen leading to lysis. Additionally, one of the products of the cleavage of C3, C3b, acts as an opsonin leading to the recognition and engulfment of the pathogen by phagocytic cells. Other complement proteins including C3a, C4a, and C5a act as inflammatory agents increasing vascular permeability and recruiting phagocytes (Vonarbourg et al., 2006).  1.5.3.3 Strategies for Improving Hemocompatibility Several strategies have been developed to alter the structure of polymeric surfaces to prevent thrombosis and recognition by the complement system, thus producing  78 hemocompatible polymers. The grafting of hydrophilic polymers to blood contacting surfaces has been shown to be effective in the prevention of protein adsorption and the initiation of the coagulation cascade and recognition by the complement system. Several coating materials have been investigated, including polysaccharides such as dextran, heparin and chitosan. However, PEG is the most widely used coating material (Lemarchand et al., 2004; Vonarbourg et al., 2006). Due to its highly hydrophilic nature and ability to hydrogen bond to water molecules, PEG acts to shield the surface of nanoparticles by forming a dense hydrophilic cloud over the surface. Several factors including the molecular weight, surface density and special conformation and flexibility of PEG chains have been shown to play a part in the ability of polymers to effectively decrease the binding of proteins to the coated substrate (Vonarbourg et al., 2006). The first attempts at surface PEGylation included adsorbing surfactants containing PEG blocks to the surface of the nanoparticles. Materials including Pluronics and PEG- PDLLA diblock copolymers were investigated, in which the hydrophobic blocks of the copolymer adsorb to the surface of the nanoparticle, with the hydrophilic PEG chains becoming hydrated and protruding from the surface (Gref et al., 1995; Mueller and Wallis, 1993; Troester and Kreuter, 1992). Due to desorption of the surfactant, it was demonstrated that uptake by the RES was inevitable and methods of retaining PEG on the nanoparticle surface were explored (Gref et al., 1995). Covalently bonding PEG to the surface, or forming the nanoparticles entirely from PEG diblock copolymers, improves the stability of the coating, resulting in prolonged circulation (Gref et al., 1994). Methods of preventing unwanted coagulation include the stimulation of thrombin inhibition or the stimulation of fibinolysis by conjugating heparin or plasminogen to polymer surfaces,  79 respectively. Heparin coatings are also currently used to prevent the complement activation of PMMA intraocular lenses, as well as in the extracorporeal circuits in cardiopulmonary bypass and renal dialysis (Benoit, 2005).  80  1.6 THESIS RATIONALE AND RESEARCH OBJECTIVES 1.6.1 Rationale The hydrophobic anticancer drug, PTX, is commercially formulated in a mixture of Cremophor! EL and ethanol. However, this excipient has been implicated in a number of adverse effects (Dorr, 1994), leading to considerable interest in the development of new biodegradable amphiphilic block copolymer formulations for PTX. Ideally, this formulation would be non-toxic and increase the aqueous solubility of PTX. Additionally, this system should provide controlled release of the drug and circulate in the blood for sufficient time to allow the passive targeting of tumors. Several amphiphilic diblock copolymer formulations are effective at solubilizing PTX (Deng et al., 2005; Zhang et al., 1996); however, as demonstrated by in vitro plasma distribution studies (Ramaswamy et al., 1997) and in vivo pharmacokinetic and biodistribution studies (Burt et al., 1999; Le Garrec et al., 2004), often PTX does not remain within the nanoparticles and thus, is not subject to increased blood circulation time and passive targeting of tumors. It was the overall goal of this project to develop a non-toxic, amphiphilic block copolymer nanoparticulate delivery system that would not only increase the aqueous solubility of PTX, but also retain the drug in plasma. Such a formulation should provide prolonged circulation to enable the passive targeting of the drug delivery system to tumor sites.  MePEG-b-PCL diblock copolymers were selected for formulating the nanoparticulate drug delivery system. MePEG-b-PCL has been used for the solubilization of several  81 hydrophobic drugs including, rapamycin (Forrest et al., 2006), indomethacin (Kim et al., 2001), and cyclosporine A (Aliabadi et al., 2005). We considered this copolymer to be an appropriate candidate for drug delivery applications as PCL is biocompatible and has a history of use in biomedical applications including implants and sutures (Nair and Laurencin, 2007). Its high degree of hydrophobicity and prolonged degradation time make PCL a good choice for controlled drug delivery. Furthermore, our group has a particular interest in drug formulations based on MePEG-b-PCL amphiphilic diblock copolymers, based on an earlier observation that some members of a series of short block length MePEG-b-PCL copolymers reduced the efflux transport of some hydrophobic drug substrates by Pgp, a protein that is implicated in mechanisms of drug resistance in tumor cells and decreased drug absorption through Pgp overexpressing epithelia (Zastre et al., 2002).  It has been shown that the physicochemical properties of amphiphilic block copolymers, as well as the nanoparticulate delivery systems they form, are highly dependent on the relative block lengths of the copolymer (Riley et al., 2001). Therefore, to develop an amphiphilic block copolymer nanoparticulate drug delivery system that will effectively solubilize and retain PTX, it is essential to understand how the composition of the block copolymer relates to the performance of the nanoparticles. In order to do so, a series of amphiphilic block copolymers with varying hydrophilic and hydrophobic block lengths were synthesized. The physicochemical properties including, thermodynamic stability, phase behavior, particle size and drug solubilization were assessed to aid in the selection  82 of copolymers for further evaluation of hemocompatibility and in vitro transport and partitioning of PTX in plasma components.  1.6.2 Research Objectives 1) To synthesize and characterize a series of amphiphilic diblock copolymers based on methoxy poly(ethylene glycol)-block-poly(!-caprolactone) with varying hydrophilic and hydrophobic block lengths  2) To investigate the effects of changes in hydrophilic (MePEG) and hydrophobic (PCL) block lengths on the physicochemical properties of MePEG-b-PCL diblock copolymers and the nanoparticulate systems formed.  3) To develop empirical relationships between the extent of micellar solubilization of several hydrophobic drugs and both the compatibility between the drug and core-forming block (PCL), as determined by the calculated "sp value, and the PCL block length.  4) To differentiate between MePEG-b-PCL micelles and nanospheres with respect to their physicochemical properties and relate these differences to changes in hemocompatibility, and in vitro plasma distribution of paclitaxel.  83  1.7 REFERENCES  Ahmed F and Discher D E. 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Biomaterials 26 (2005) 2121-2128.  102  CHAPTER 2 SYNTHESIS AND CHARACTERIZATION OF METHOXY POLY(ETHYLENE GLYCOL)-BLOCK-POLY(!-CAPROLACTONE) DIBLOCK COPOLYMERS 1   2.1 INTRODUCTION Various aliphatic polyesters such as poly(lactic acid) (PLA), poly(d,l-lactic-co-glycolic acid) (PLGA) and poly(!-caprolactone) (PCL) have been covalently bonded with a hydrophilic PEG segment to produce diblock or triblock copolymer structures (Allen et al., 1998; Burt et al., 1999; Hagan et al., 1996; Shin et al., 1998; Yoo and Park, 2001). Most often these copolymers are used for the production of nanoparticles due to their ability to self-assemble forming core-shell structures such as micelles and nanospheres or vesicles such as polymersomes. Other delivery systems such as thermoreversible gels have been explored (Jeong et al., 2000).  Block copolymers of PEG and polyesters may be synthesized by a variety of methods. It is possible to use “living polymerization” techniques to polymerize ethylene oxide resulting in poly(ethylene oxide) as the hydrophilic block leaving the terminal end of PEO with a hydroxyl group. Most commonly, this is achieved by using sequential anionic polymerization in which an anionic initiator such as diphenyl methyl potassium is used to ring open the monomer, ethylene oxide, which in turn is allowed to propagate until completion of the reaction. The lactone is then added to the mixture and is polymerized  1 A version of this chapter has been published. Letchford K et al., (2004) Synthesis and micellar characterization of short block length methoxy poly(ethylene glycol)-block- poly(caprolactone)diblock copolymers. Colloids Surf., B 35(2) 81-91.  103 by the living end of the poly(ethylene oxide). This type of reaction has been used extensively by Eisenberg’s group for the synthesis of PEO-b-PCL diblock copolymers (Allen et al., 1998). A much simpler method involves the use of a preformed PEG block such as !-methoxy-"-hydoxyl-PEG (MePEG). In this reaction the MePEG acts as a macroinitiator for the ring opening of the lactone via nucleophilic attack of the ester bond by the terminal hydroxyl group of MePEG. These reactions are usually carried out in the presence of stannous octoate as a catalyst at elevated temperatures; however, some groups have demonstrated that the reaction may proceed successfully by cationic ring opening using hydrogen chloride diethyl ether (Lee et al., 2005), or alternatively in the absence of a catalyst (Lin et al., 2003).  Recently, amphiphilic copolymers consisting of PEG and PCL have received a great deal of attention. Due to its proven history of biocompatibility of use in medical devices such as sutures, stents, prosthetics and as a carrier for hydrophobic drugs (Pitt, 1990; Winternitz et al., 1996), PCL is an excellent choice for the use as the hydrophobic block of amphiphilic copolymers. PCL is a synthetic, semicrystalline, biodegradable polyester that is considerably more hydrophobic than other aliphatic polyesters allowing the polymer to degrade slowly making it a good candidate for applications of controlled release (Pitt, 1990; Pitt et al., 1981; Woodward et al., 1985). Block copolymers of PEG and PCL with PEG MW ranging from 2000-5000 g/mol and 14-50 PCL repeat units have been evaluated as micellar carriers of hydrophobic drugs including dihydrotestosterone (Allen et al., 2000), indomethacin (Shin et al., 1998) and paclitaxel (Kim and Lee, 2001). We have demonstrated that a novel series of MePEG-b-PCL diblock copolymers  104 consisting of MePEG with a molecular weight of 750 g/mol and PCL blocks lengths ranging from 2-10 repeat units are effective at modulating P-glycoprotein (P-gp) mediated efflux of various P-gp substrates in Caco-2 cells and multi drug resistant cancer cells (Zastre et al., 2002). Inhibition of this efflux pump through the use of polymeric amphiphilic agents may be important in the successful delivery drugs through epithelial cells expressing P-gp, such as intestinal and blood brain barrier tissues (Batrakova et al., 1999) as well as overcoming multi drug resistance in cancer cells (Alakhov et al., 1996).  The objective of this work was to characterize the synthesis and physicochemical properties of a homologous series of MePEG-b-PCL copolymers possessing 17-114 repeat units of MePEG and 2-104 repeat units of PCL, in order to understand how the properties of the diblock copolymers varied with changes to block lengths. In subsequent chapters of this thesis, diblock copolymers were selected from this series with significantly different block lengths, to fabricate nanoparticles with distinctly different core-shell properties and drug delivery performance characteristics.  2.2 EXPERIMENTAL 2.2.1 Materials Methoxy poly(ethylene glycol) (MePEG) with molecular weights of 750, 2000 or 5000 g/mol (MePEG) and !-caprolactone (CL) were purchased from Fluka (Oakville, ON). Stannous octoate and pyrene were obtained from Sigma-Aldrich (Oakville, ON) and were used as supplied without further purification. Chloroform and acetone were HPLC grade and used as supplied by Fisher Scientific (Ottawa, ON). Deuterated chloroform was  105 obtained from Cambridge Isotopes (Andover, MA). Poly(ethylene glycol) molecular weight standards were purchases from Polymer Laboratories (Amherst, MA). Platinum/palladium wire (0.22 mm diameter) from Canemco (Canton de Gore, PQ) was used for shadowing for the TEM studies.  2.2.2 Synthesis of Methoxy Poly(ethylene glycol)-Block-Poly(!-caprolactone) Diblock Copolymers MePEG with a MW of 750, 2000 or 5000 g/mol was combined with CL in varying weight ratios to control the final MW of the copolymer. The reagents, totaling 50 g, along with a teflon coated stir bar were placed in a round bottom flask sealed with a ground glass stopper and immersed in a heavy mineral oil bath heated to 140°C. The reagents were mixed for 30 minutes to produce a homogenous liquid, at which time 0.15 ml of stannous octoate were added to the flask. The polymerization reaction was allowed to proceed for 24 hours. Cooling the polymer to room temperature terminated the reaction. No further purification steps were performed as NMR and GPC analysis demonstrated negligible amounts of unreacted monomer and homopolymer.  2.2.3 Nuclear Magnetic Resonance Spectroscopy (NMR) Copolymer samples were dissolved in deuterated chloroform at a concentration of 10% w/v for analysis by 1 H NMR using a Bruker (Billerica, MA) 400 MHz NMR instrument. Peak positions and areas were analyzed to determine the degree of polymerization using  106 MestRe-C 2.3a software (Mestrelab Research, Santiago de Compostela, Spain) and comparing integrals of peaks from the PCL and PEG blocks.  2.2.4 Gel Permeation Chromatography (GPC) Copolymer molecular weights were determined by gel permeation chromatography (GPC) against poly(ethylene glycol) standards in the range of 670 – 22800 g/mol using a Waters (Milford, MA) GPC system. Samples were injected using a Waters model 717 plus autosampler. Chloroform with a flow rate of 1 ml/min was used as the mobile phase and separation was achieved through two Waters Styragel columns (HR 3 and HR 1) connected in series. Detection was by a Waters model 2410 refractive index detector with a cell temperature of 40°C. Residual unreacted !-caprolactone and consumption of !- caprolactone were determined using the same system and !-caprolactone standard solutions.  2.2.5 Critical Micelle Concentration (CMC) and Pyrene Partition Equilibrium Constant (Kv) The critical micelle concentration (CMC) and partition equilibrium constants (Kv) were determined by a steady state pyrene fluorescence method (Lee et al., 1999; Lele and Leroux, 2002; Wilhelm et al., 1991). A 1.5 x 10 -5  M solution of pyrene in acetone was aliquoted (320 µL) into a series of amber vials and the acetone was evaporated under a stream of nitrogen. To the vials were added 8 ml of aqueous copolymer solution ranging from 0.0001 to 10 g/L in PBS (pH =7.4) to give a final pyrene concentration of 6.0 x10 -7   107 M. Samples were incubated for 24 hours at 37°C in the dark with stirring. Samples (2 ml) were pipetted into Sarstedt (Montreal, PQ) 10 x 10 x 48 mm acryl disposable cuvettes and the excitation spectra were recorded at 37°C on a Shimadzu (Columbia, MD) RF 540 spectrofluorometer with an emission wavelength of 390 nm and excitation and emission slit widths of 2 nm. An additional set of studies were carried out at 30°C to determine the CMC of MePEG17-b-PCL2 at this temperature in order to complete the phase diagram for this polymer.  2.2.6 Hydrodynamic Diameter Light scattering measurements were carried out on a Malvern (Malvern, UK) 3000HS Zetasizer  with a He-Ne laser (532 nm) and 90° collecting optics. Copolymer samples at a concentration of either 1% w/v (or 0.1% w/v for MePEG17-b-PCL10) in PBS were prepared and filtered using a Millipore (Billerica, MA) 0.22 µm filter  prior to measurement. All measurements were carried out at 37°C. Data were analyzed using CONTIN algorithms provided by Malvern ZetaSizer software. 2.2.7 Aqueous Solution Phase Behavior and Micelle Morphology of Short Block Length Methoxy Poly(ethylene glycol)-Block-Poly(!-caprolactone) Diblock Copolymers Copolymer phase behaviour with respect to temperature was determined by heating or cooling polymer samples dissolved in phosphate buffered saline (PBS) ranging in concentration in a Fisher Scientific (Ottawa, ON) Isotemp 210 water bath. Samples were heated at 40°C until the polymer dissolved then the temperature was adjusted by ± 1°C  108 increments over a temperature range of 22-60°C. Sample temperatures below 22°C were achieved using a model EK12 Haake (Thermo Scientific, Newington, NH) chiller connected to a water bath with a model C1 Haake heater/circulator. Samples were allowed to equilibrate for 10 minutes after the desired temperature was reached. Cloud points and Krafft points were taken to be the temperature at which solutions became turbid, determined by visual inspection.  Transmission electron microscopy (TEM) was used to observe the morphology and measure the size of micelles. Samples of 5 µL of dilute diblock copolymer solutions (0.1 – 0.01 % w/v) were pipetted on to copper EM grids precoated with Formvar! and carbon. Solutions were allowed to stand for a few minutes before being wicked off with hardened filter paper. The grids were dried overnight at room temperature and pressure before shadowing with Pt/Pd alloy at approximately a 30° angle. TEM was carried out on a Hitachi (Pleasanton, CA) H7600 transmission electron microscope operating at an accelerating voltage of 80 kV.  2.3 RESULTS 2.3.1 Synthesis and Characterization of MePEG-b-PCL Diblock Copolymers In this study we synthesized a series of MePEG-b-PCL diblock copolymers by a ring opening mechanism of CL using the hydroxyl group of MePEG as the initiator and stannous octoate as a catalyst (figure 2.1). The molecular weight of MePEG used in the synthesis was 750 g/mol, 2000 g/mol or 5000 g/mol and varying the feed ratio of MePEG to CL in the reaction mixture controlled the final amount of PCL in the copolymer (Table  109 2.1). GPC was used to estimate the MW of the synthesized copolymers and the amount of unreacted !-caprolactone in the final product. Samples based on MePEG with a MW of 750 g/mol and a 60:40 (MePEG : !-caprolactone) feed stock were prepared and allowed to react for times ranging from 0 – 48 hours. The MW and percentage of unreacted !- caprolactone were determined. Chromatograms of these samples are shown in figure 2.2. By 8 hours of reaction, the CL peak at 17.4 minutes was undetectable, while the MePEG peak, originally at 13.7 minutes, increased in size and shifted to an earlier retention time of 13.1 minutes. A plot of MW and % unreacted !-caprolactone as a function of reaction time is shown in figure 2.3. The molecular weight of the synthesized copolymer and the percentage of unreacted CL began to plateau by 8 hours of reaction as shown in figure 2.3. All samples used for further studies were reacted for a minimum of 8 hours. The majority of the synthesized diblock copolymers had a narrow MW distribution, indicated by a polydispersity index (PDI) close to 1 as determined by GPC, and MW values which were similar to the theoretical MW (Table 2.1). The only exception was the copolymer with a 30:70 feed ratio of MePEG 5000 g/mol to CL, which had a slightly higher PDI than the rest of the diblocks and a GPC MW that deviated from the theoretical and NMR MWs.  110     CH3O CH2CH2O n H OC O Sn(Oct)2 140oC CH3O CH2CH2O n CCH2(CH2)3CH2O O m H Methoxy poly(ethylene glycol) Caprolactone MePEG-b-PCL    n is the degree of polymerization (block length) of MePEG. n was one of: 17 repeat units – MePEG MW 750 g/mol  44 repeat units – MePEG MW 2000 g/mol  114 repeat units – MePEG MW 5000 g/mol  m is the degree of polymerization of PCL and ranged from 2 – 104 repeat units                Figure 2.1 Polymerization reaction of MePEG-b-PCL.   111 Table 2.1 Synthesis and characterization data for methoxy poly(ethylene glycol)-block-poly(!-caprolactone) diblock copolymers MePEG MW (g/mol) Feed Ratio a M/Itheo M/INMR DPn b MWtheo c (g/mol) MWNMR d (g/mol) MWGPC e (g/mol) PDI f CMC g (mol/l) Kv h Dh i (nm) 750 80:20 1.64 1.75 MePEG17-b-PCL2 939 947 980 1.13 1.89x10 -3 6.99x10 3  12.2±1.0 750 70:30 2.82 2.83 MePEG17-b-PCL3 1071 1073 1096 1.19 1.11x10 -4 3.88x10 4  13.5±0.3 750 60:40 4.38 3.85 MePEG17-b-PCL4 1250 1214 1287 1.19 5.96x10 -5  4.15x10 4  14.5±0.1 750 40:60 9.85 10.0 MePEG17-b-PCL10 1875 1890 1813 1.07 1.79x10 -6 2.22x10 5 19.5±1.2 2000 70:30 7.51 7.47 MePEG44-b-PCL7 2857 2852 2594 1.08 7.36x10 -6 6.81x10 4  14.7±0.2 2000 60:40 11.68 11.82 MePEG44-b-PCL12 3333 3320 2825 1.24 1.69x10 -6 1.34x10 5  22.3±0.7 5000 70:30 18.77 19.0 MePEG114-b-PCL19 7143 7166 7044 1.18 6.29x10 -7 2.80x10 5  45.3±1.3 5000 30:70 102.2 103.6 MePEG114-b-PCL104 16651 16651 12238 1.47 * * * a  Feed weight ratio of MePEG : caprolactone in reaction mixture. b  Degree of polymerization (DPn) of MePEG: MWMePEG/44, DPn of PCL: rounded off value determined by NMR. c  Theoretical MW determined by feed ratio. d  MW calculated by DPn according to NMR. e  MW determined by GPC. f  Polydispersity index determined by GPC. PDI = Mw/Mn. g  Critical micelle concentration determined by pyrene fluorescence excitation spectra at 37°C. h  Partition coefficient of pyrene between the core forming block and aqueous phase determined by pyrene fluorescence excitation spectra at 37°C. i  Micelle hydrodynamic diameter determined by dynamic light scattering at 37°C. *Due to the poor water solubility of this copolymer CMC, Kv, and Dh values were not determined.  112                              Figure 2.2 GPC chromatograms of the polymerization mixture with a feed ratio of 60:40 of MePEG 750 g/mol and CL. Reaction times are indicated on the left. At 0 hrs, peak A represents MePEG and peak B is CL. 0 h 4 h 8 h 12 h 24 h 48 h A B Retention Time (min)  113              Figure 2.3 The effect of reaction time on the number average molecular weight (Mn) (!) of MePEG-b-PCL and weight percent of unreacted caprolactone during the synthesis of MePEG17-b-PCL4 ("). Data points and error bars represent the mean ± S.D. (n =3).  114 The 1 H NMR spectra were used to determine the degree of polymerization of each synthesized copolymers. Comparisons were made of the integrated peak area of the multiplet at 3.55 ppm, assigned to the methylene groups of MePEG, and the sum of the integrated peak areas of the multiplets at 1.5 ppm and 1.3 ppm respectively, assigned to methylene protons in the caprolactone repeat unit (Zastre et al., 2002) (Figure 2.4 and Table 2.1). The number of MePEG repeat units was calculated by dividing the molecular weight of MePEG by 44 (the MW of the repeat unit), giving approximately 17, 44 and 114 for MePEGs with molecular weights of 750 g/mol, 2000 g/mol and 5000 g/mol, respectively. Theoretical molecular weights were calculated from the feed ratios of MePEG and caprolactone whereas the NMR molecular weights were calculated from the degree of polymerization values determined by NMR. Neither the NMR spectra nor the GPC chromatograms showed any evidence of additional compounds such as unreacted monomer or side reaction products, therefore, no further purification steps were performed. The nomenclature used for the diblock copolymers displays the degree of polymerization of each component as a subscript behind each block abbreviation, rounded to the nearest whole number. For example, the copolymer MePEG114-b-PCL19 has 114 repeat units of MePEG and 19 repeat units of caprolactone.  115              Figure 2.4 A representative 1 H NMR spectrum of a MePEG-b-PCL diblock copolymer using a Bruker 400 MHz spectrometer.  116 When exposed to an aqueous environment, amphiphilic diblock copolymers of MePEG- b-PCL are capable of forming micelles when present at or above a specific concentration termed the critical micelle concentration (CMC). The CMC of a copolymer is an indication of the thermodynamic stability of a micellar formulation upon dilution. In this study the CMC of all the copolymers with the exception of MePEG114-b-PCL104 were determined using the fluorescence probe pyrene (Lee et al., 1999; Lele and Leroux, 2002; Wilhelm et al., 1991). Due to its poor aqueous solubility, MePEG114-b-PCL104 could not be dissolved in water, therefore the CMC could not be determined. The (0,0) band in the excitation spectra of pyrene shifted from 333 nm to 336 nm as the probe partitioned from the aqueous environment to the hydrophobic core of the micelle (figure 2.5). The ratio of the fluorescence intensity at 336 nm to the fluorescence intensity at 333 nm (I336nm/I333nm) was calculated for each polymer concentration and plotted versus the logarithm of polymer concentration. Figure 2.6 shows a representative plot of the intensity ratios (I336nm/I333nm) versus the logarithm of diblock copolymer concentration. In this figure, it can be seen that at low copolymer concentrations, the intensity ratio is relatively constant until a critical concentration is reached, at which time there is a rapid increase in the intensity ratio. The rapid increase in intensity ratio is attributed to the probe incorporating into the hydrophobic core of the micelles and provides an estimate of the CMC. The CMC was determined from the intersection of the two straight lines (horizontal line of I336nm/I333nm with almost constant values and the diagonal line with an increase in the I336nm/I333nm values) on a plot of the ratio I336nm/I333nm versus log polymer concentration (figure 2.6). Figure 2.7 shows plots of intensity ratio versus logarithm of copolymer concentration for the series of MePEG-b-PCL copolymers synthesized. In comparing the  117 CMC values between copolymers, it was apparent that there was an inverse correlation between the CMC and the PCL block length, with CMC values ranging from 1.89 x 10 -3  mol/L to 6.29 x 10 -7  mol/L as the PCL block length increased from 2 to 19 repeat units, respectively (Table 2.1). In a previous study, we determined the CMC of the MePEG 750 g/mol copolymer series by a fluorescence method using 1,6-diphenyl-1,3,5-hexatriene (DPH) (Zastre et al., 2002). These CMCs were higher than values determined in this study. The CMC of MePEG17-b-PCL2 at 30°C was determined to be slightly higher (2.69 x 10 -3  mol/L) than that at 37°C.  118          Figure 2.5 Fluorescence excitation spectra of pyrene solubilized in aqueous solutions of MePEG17-b-PCL4. Emission wavelength 390 nm and 37°C.  119 Measurements of the partition equilibrium constant (Kv) of pyrene in the micellar solutions were used to estimate the hydrophobicity of the micelle core The method used was similar to that of Wilhelm et al. (Wilhelm et al., 1991) who proposed the equations:  ! [Py]m [Py]w = (F "Fmin ) (Fmax "F)   (2.1) and ! [Py]m [Py]w = Kv"PCLC 1000#   (2.2) therefore, ! (F "F min) (F max"F) = Kv#PCLC 1000$   (2.3)  where [Py]m and [Py]w are the concentration of pyrene in the micellar and aqueous phases, respectively. Fmin and Fmax are the average I336nm/I333nm ratios in the horizontal regions of the I336nm/I333nm vs. log polymer concentration plots. F is the average I336nm/I333nm ratios for the intermediate copolymer concentrations (figure 2.6). !PCL is the weight fraction of PCL in the block copolymer. C is the copolymer concentration and " is the density of PCL in the core, which was approximated as the density of bulk PCL (1.146g/mL) (Lee et al., 1999). Kv was obtained from the slope of a plot of (F-Fmin)/(Fmax- F) vs [polymer] (figure 2.8). The Kv values for all the copolymers were determined by this method with the exception of MePEG114-b-PCL104, which was too hydrophobic to be dispersed in the aqueous phase. The partition coefficient of pyrene between the core  120 forming block and water was found to increase from 6.99 x 10 3  to 2.80 x 10 5  as the PCL block length increased from 2 to 19 repeat units (Table 2.1).  121            Figure 2.6 Representative plot of I336nm/I333nm as a function of copolymer concentration for the determination of CMC and Kv. Fmax Fmin  122                     Figure 2.7 Plots of pyrene fluorescence intensity ratio I336nm/I333nm at 37°C as a function of polymer concentration for the determination of CMC. MePEG17-b-PCL2 (!),MePEG17-b-PCL3 (!),MePEG17-b-PCL4 (!), MePEG17-b-PCL10 ("), MePEG44-b-PCL7 (!), MePEG44-b-PCL12 (") and MePEG114-b-PCL19 (#). Data points and error bars represent the mean ± S.D. (n =3).  123                            Figure 2.8 A representative plot of (F-Fmin)/Fmax-F) vs. polymer concentration for the determination of the pyrene partition coefficient (Kv) for the copolymer MePEG17-b-PCL10.  124 The average micelle diameters determined by dynamic light scattering ranged from 12 – 45 nm. There was a direct correlation between the hydrodynamic diameter and the PCL block length (Table 2.1). Micelle size distributions were found to be monomodal.  2.3.2 Aqueous Solution Phase Behaviour and Micelle Morphology of Short Block Length MePEG-b-PCL Diblock Copolymers Solutions of the diblock copolymers based on MePEG with a molecular weight of 750 g/mol were found to exhibit temperature dependent phase behavior. At relatively low temperatures, the copolymer solutions were cloudy and eventually precipitated. As the temperature was increased, the solutions became clear until a temperature was reached above which the solutions once again became cloudy (figure 2.9). Temperature versus concentration phase diagrams were constructed in order to determine the range which the MePEG 750 g/mol copolymers remained in solution (figure 2.10). Polymer solubility in aqueous media existed between two boundaries, the lower being the Krafft point and the upper being the cloud point. MePEG17-b-PCL4 displayed a constant Krafft point of approximately 30°C for copolymer concentrations between 0.1-100 g/L. MePEG17-b- PCL10 was least soluble and displayed a constant Krafft point of 31°C over the 0.1-2 g/L concentration range. Above 2 g/L MePEG17-b-PCL10 was determined to be insoluble due to the presence of precipitate at all temperatures. MePEG17-b-PCL2 had a constant Krafft point of 27°C over the 2-100 g/L range. Below 2 g/L the samples precipitated at progressively lower temperatures. At a concentration of 0.5 g/L precipitation was not observed at 4°C, the minimum temperature tested. All the MePEG-b-PCL copolymer solutions tested displayed cloud points. MePEG17-b-PCL2 showed a constant cloud point  125 of 53°C between 20-100 g/L and the cloud point increased at lower concentrations. MePEG17-b-PCL4 and MePEG17-b-PCL10 had constant cloud points of approximately 54°C and 51°C between the concentration ranges of 0.1-100g/L and 0.5-2 g/L, respectively and the cloud points also increased at lower concentrations.  126                        Figure 2.9 Photographs of 20 g/L aqueous solutions of MePEG17-b-PCL2 at 4°C (below the Krafft point), 37°C (micellar solution), and 60°C (above the cloud point). Temp = 4°C Temp = 37°C Temp = 60°C  127    Figure 2.10 Phase diagrams of MePEG-b-PCL diblock copolymers with a MePEG MW of 750 g/mol. (A) MePEG17-b-PCL10, (B) MePEG17-b-PCL4, (C) MePEG17-b-PCL2. (!) Cloud point, (") Krafft point, and (!) CMC. Data points and error bars represent the mean ± S.D. (n =3). A B C  128 Figure 2.11 is a series of transmission electron micrographs of micelles formed by the MePEG 750 g/mol series of copolymers. The micelles appeared as dark spheres with a light triangular tail due to the shadowing with Pt/Pd alloy. From this micrograph it can be seen that micelles formed are spherical in nature. MePEG17-b-PCL10 micelles are approximately 40 nm in diameter whereas those of MePEG17-b-PCL4 and MePEG17-b- PCL2 are approximately 20 nm in diameter.  129           Figure 2.11 Transmission electron micrographs of MePEG-b-PCL micelles. Shadowed at a 30° angle with Pt/Pd alloy. Accelerating voltage 80kV. (A) MePEG17- b-PCL2 (B) MePEG17-b-PCL4, (C) MePEG17-b-PCL10. A B C  130  2.4 DISCUSSION MePEG-b-PCL diblock copolymers were synthesized by a ring opening polymerization of !-caprolactone using stannous octoate as a catalyst and the hydroxyl group of MePEG as the initiator. The degree of polymerization could be controlled by varying the ratio of !-caprolactone to MePEG in the feed stock as evidenced by the small deviation between the theoretical monomer to initiator ratios (M/Itheo) and those determined by NMR (M/INMR) (Table 2.1). Reactions proceeded to a high degree of completion resulting in undetectable levels of unreacted !-caprolactone determined by GPC after 8 hours of reaction.  It has been well documented that when amphiphilic diblock copolymers are added to an aqueous environment at, or above, the CMC, they self-assemble to form micelles (Alexandridis et al., 1994; Gaucher et al., 2005; Jones and Leroux, 1999; Kabanov and Alakhov, 2002; Kwon and Kataoka, 1995). The CMC provides an indication of the thermodynamic stability of the micelles or the minimum concentration at which these nanoparticles will stay self-assembled. The primary driving force behind this aggregation is the increase in entropy of the surrounding water molecules upon the aggregation of the copolymers (Alexandridis and Hatton, 1995). In this study, a hydrophobic fluorescent probe, pyrene, was used to detect the CMC. Pyrene displays a characteristic shift in the fluorescent excitation spectrum from 333 nm to 336 nm upon partitioning into a hydrophobic environment such as the PCL core of the micelles. Similar to previous reports, we found an inverse relationship between the hydrophobic block length and the  131 CMC (Kang and Leroux, 2004; Kim et al., 2000; Zastre et al., 2002). This trend can be explained thermodynamically as follows. As the hydrophobic block length of the copolymer increases, so does the amount of unfavorable interactions between this block and the surrounding water molecules, causing an increase in the structuring of the water surrounding the hydrophobic blocks of the copolymer, decreasing the entropy of the system (Alexandridis and Hatton, 1995). As the number of molecules of copolymer in solution are increased, the entropy, and hence the free energy of the system increases until a threshold free energy level is obtained, at which point micelles are formed in order to reduce the amount of free energy in the system (Alexandridis and Hatton, 1995). Since copolymers with longer hydrophobic blocks have more unfavorable interactions per molecule than those with shorter hydrophobic blocks, the threshold free energy and therefore the CMC occurs at a lower polymer concentration. Interestingly, even with the large increases in the MePEG block length, which can cause increases in CMC, the modest increases in PCL block length still produced a decrease in CMC, confirming previous reports that the hydrophobic block length is the primary factor determining the CMC (Alexandridis et al., 1994).  When considering the development of a micellar formulation of a hydrophobic compound it is important to consider whether the drug payload will remain in the micelle structure upon injection in the body and during subsequent dilution in the bloodstream. The CMC of the copolymer indicates at which point the micelles will disassemble upon dilution. It is likely that not all copolymers used in this work would make good candidates for in vivo application, as dilution would lead to disassembly and loss of the  132 loaded drug, due to relatively high CMC values (e.g. MePEG17-b-PCL2). However, even though a formulation may possess a low CMC, this factor alone is not sufficient to predict that the drug will remain within the micelle. Our group has shown, in the case of paclitaxel loaded MePEG-b-PDLLA micelles, that the drug rapidly dissociated from the copolymer into the blood regardless of the low CMC (Burt et al., 1999). Interactions with various blood components are responsible for this dissociation as will be discussed in a later chapter.  The CMCs of the MePEG-b-PCL copolymers in this study were comparable to other copolymers utilizing PCL as the hydrophobic block. Allen et al. determined the CMC of PEO44-b-PCL21 to be 2.8x10 -7  M and Shin et al. determined the CMC of MePEG114-b- PCL35 to be 3.53x10 -7  M (Allen et al., 1999; Shin et al., 1998). When compared to copolymers composed of polyester hydrophobic blocks other than PCL, our copolymers possessed low CMCs, despite their relatively short hydrophobic block lengths, most likely due to the greater degree of lipophilicity of PCL. Hagan et al. found the CMC of MePEG44-b-PLA13 and MePEG114-b-PLA11 to be 1.17x10 -5  M and 6.0 x 10 -6  M, respectively (Hagan et al., 1996) whereas Kabanov reported the CMC for Pluronics! to range from 2.27 x 10 -6  M for PEO5-b-PPO34-PEO5 (Pluronic ! L121) to 7.65 x 10 -3  M for PEO11-b-PPO8-PEO11 (Pluronic ! L35) (Kabanov et al., 2002). Kwon et al. reported the CMC of the series of block copolymers MePEG110-b-PBLA9, MePEG110-b-PBLA19, MePEG270-b-PBLA20 to be 1.43x10 -6 , 5.49x10 -7 , 6.25x10 -7  M, respectively (Kwon et al., 1993).   133 As previously mentioned, the CMC values for the MePEG 750 g/mol copolymer series found by the pyrene method were lower than those determined using DPH fluorescence (Zastre et al., 2004). This has been noted by other groups (Kabanov et al., 1995) and is probably due to the ability of DPH to form supermolecular complexes in aqueous solutions that can undergo time dependent fluorescence changes during incorporation in to micelles leading to erroneous CMC values (Kabanov et al., 1995). The pyrene data are believed to be more accurate (Kabanov et al., 1995) and most CMC studies have utilized this method.  The CMC of MePEG17-b-PCL2 decreased as the temperature increased from 30°C to 37°C. This trend is well documented for non-ionic surfactants and may be explained by a dehydration of unimers at elevated temperatures leading to an increase in hydrophobicity and a subsequent decrease in the CMC (Alexandridis et al., 1994; Attwood and Florence, 1982).  Due to its insolubility in aqueous media, neither the CMC nor Kv were determined for MePEG114-b-PCL104. This polymer was used in a later chapter to fabricate nanospheres and shown to possess distinctly different physicochemical properties compared to the micelles described in this chapter.  As seen in table 2.1, the partition equilibrium constants increased as the PCL block length increased, a trend documented for other copolymers (Kozlov et al., 2000; Wilhelm et al., 1991). When compared to partition equilibrium constants of other block copolymers,  134 these MePEG-b-PCL diblock copolymers have relatively high Kv values. Pluronic ! P85 and F108 were reported to have pyrene Kv values of 3260 and 1640, respectively (Kabanov et al., 1995). Kv values in the range of 3 x 10 5  were found for PEO-b- poly(styrene) (Zhao and Winnik, 1990). Kwon et al. found the Kv for pyrene in PEO-b- poly(" benzyl-L-aspartate) to be of the order of 104 (Kwon et al., 1994) and copolymers consisting of PCL hydrophobic cores display a range of Kv ranging from 10 2  to 5.88x10 5  (Allen et al., 1998; Lee et al., 1999). A factor that may account for the high Kv values found for our diblock copolymers is that the Kv determination was conducted at 37°C whereas values obtained for the other copolymer systems, with the exception of the pluronics (Kabanov et al., 1995), were conducted at either 20°C (Allen et al., 1998) or 25°C (Kwon et al., 1994; Lee et al., 1999; Zhao et al., 1990). It has been found that upon increasing the temperature from 20°C to 37°C the Kv of pyrene in pluronic micelles increased by an order of magnitude (Kabanov et al., 1995). It was proposed that the reason for this increase in Kv was due to the dehydration of the core-forming blocks causing an elevation of the micellar core hydrophobicity.  Upon experimentation with the MePEG 750 g/mol copolymer series, it was noticed that they exhibited temperature dependent phase transitions. To further investigate this behavior and characterize the temperature and concentration ranges that these diblock copolymers could be formulated, the phase behaviour of the copolymers in aqueous media was investigated and temperature versus concentration phase diagrams were created (figure 2.10). All the polymers existed as isotropic aqueous micellar solutions at  135 temperatures between two boundaries, the lower being the Krafft point and the upper being the cloud point.  Below the Krafft point, non-associated unimers and solid hydrated polymer exist in equilibrium. As the temperature is raised, the Krafft point is reached at which the solid hydrated polymer dissolves and is incorporated into micelles forming an isotropic phase (Shinoda et al., 1962) (Figure 2.12). Above the Krafft point, the solubility of the copolymer dramatically increases since the free unimers aggregate into micelles. Krafft point phenomena are normally observed in anionic and cationic surfactants and are rarely observed in non-ionic surfactants, since their CMC values are generally very low and their solubility does not decrease significantly with decreasing temperature and often the freezing point of water is reached before the solubility equals the CMC (Schott and Han, 1976). MePEG17-b-PCL2 had a lower Krafft point than MePEG17-b-PCL4 and MePEG17- b-PCL10. This observation may be explained by the hydrophobicity of the copolymers. Since the Kraft point is the temperature at which hydrated solid amphiphiles dissolve and form micelles, the more hydrophobic MePEG17-b-PCL4 and MePEG17-b-PCL10 require higher temperatures (and thus, a higher Krafft point) to become hydrated enough to dissolve and form a micellar solution. MePEG17-b-PCL2 is relatively hydrophilic and therefore dissolves at a lower temperature reflected by a lower Krafft point.  Above the cloud point the solutions separated into two phases, one surfactant rich and the other coexisting phase having a surfactant concentration approximately equal to the CMC (Shinoda et al., 1962). This phase separation is caused by the dehydration of PEG chains  136 upon temperature increase by either the breaking of structured water formed at low temperatures, the breaking of hydrogen bonds between water and PEG chains or changes in PEG to a less polar conformation (Mansur et al., 1997). Regardless of the mechanism, at elevated temperatures, aggregation and precipitation occur due to a decreased hydration of the PEG chains leading to an increase in micelle size above the threshold temperature and subsequent phase separation (Schott, 1969) (figure 2.12). A slight decrease in the cloud point was observed as the PCL block length increased from MePEG17-b-PCL4 to MePEG17-b-PCL10. This trend can be explained by the increase in hydrophobicity of the copolymers as the PCL block length increases. Upon dehydration of the MePEG shell of the micelles, hydrophobic-hydrophobic or van der Waal’s interactions were provided by exposed PCL cores, which promote secondary aggregation and eventual phase separation (Allen et al., 1998). Micelles with cores composed of longer PCL blocks will have more PCL chains exposed compared to shorter PCL blocks and will therefore possess a greater tendency to aggregate at lower temperatures. It was observed that the cloud points and Krafft points of these polymers did not change significantly over the concentrations tested in agreement with the findings for other non- ionic surfactants (Attwood, 1968; Schott, 1969; Schott and Han, 1976).  A fourth region can be seen on the MePEG17-b-PCL2 phase diagram bound by the CMC line and the solubility curve (figure 2.10). In this region, polymer exists as non-associated unimers. As the polymer concentration increases, given the temperature is high enough, the CMC will be reached before the solubility curve and micelles will form, increasing the solubility of the polymer. If the temperature is too low, the solubility curve is crossed  137 before the CMC is reached and the polymer precipitates forming hydrated solid polymer in equilibrium with unassociated unimers. The phase boundaries between non-associated monomer and micelles or hydrated polymer was not observed in either MePEG17-b-PCL4 or MePEG17-b-PCL10 since the CMCs for these polymers are low, making it difficult to see the change in turbidity at such low concentrations by visual observation.  138                      Figure 2.12 Schematic of the concentration (dC) and temperature (dT) dependent aqueous phase behavior of short block length MePEG-b-PCL diblock copolymers.  139 The sizes of the micelles formed were between 10 – 45 nm. When comparing the diameters of the micelles formed from diblocks with the same MePEG MW, the diameter increased with PCL block length. This increase in diameter may be attributed to an increase in the core volume. The diameters of the micelles were smaller than those of other MePEG-b-PCL micelle systems previously studied. Allen et al. found that micelles composed of PEO44-b-PCL14 or PEO44-b-PCL20 had hydrodynamic diameters of 55 nm and 62 nm, respectively (Allen et al., 1998). Shin et al. produced MePEG114–b-PCL35-100 micelles ranging in diameter from 54 nm to 130nm, again considerably larger than the micelles in this study (Shin et al., 1998). The TEM micrographs of MePEG17-b-PCL10 micelles showed a size population consisting of micelles, considerably larger than the value determined by light scattering. This size difference may be explained by secondary aggregation of several smaller micelles to form a larger mass, which may have occurred during the drying process required for electron microscopy. Micrographs of MePEG17-b- PCL2 and MePEG17-b-PCL4 also show micelles with diameters larger than those determined by light scattering but only slightly. These micelles are possibly more resistant to secondary aggregation since they possess shorter hydrophobic blocks, reducing the degree of attraction for one another. The addition of a metal coating during shadowing may also account for any size difference.  2.5 CONCLUSIONS In these studies we synthesized a series of MePEG-b-PCL diblock copolymers with varying MePEG and PCL block lengths. The length of the PCL block was controlled by the weight ratio of !-caprolactone added to the reaction mixture. All copolymers with the  140 exception of MePEG114-b-PCL104 were water soluble and capable of forming micellar solutions. The CMCs for the copolymers were inversely proportional to the PCL block length whereas the partition coefficient of pyrene and hydrodynamic diameter were found to be directly dependent on the PCL block length. Phase diagrams for aqueous solutions of MePEG-b-PCL copolymers with a MePEG MW of 750 g/mol showed characteristic temperature dependent aqueous phase behavior. 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Pharmaceutical Research 21 (2004) 1489-1497. Zhao C L and Winnik M A. Fluorescnce Probe Techniques Used to Study Micelle Formation in Water-Soluble Block Copolymers. Langmuir 6 (1990) 514-516. Zhao C L, Winnik M A, Riess G and Croucher M D. Fluorescence probe techniques used to study micelle formation in water-soluble block copolymers. Langmuir 6 (1990) 514- 516.  145  CHAPTER 3 SOLUBILIZATION OF HYDROPHOBIC DRUGS BY  METHOXY POLY(ETHYLENE GLYCOL)-BLOCK-POLY(!- CAPROLACTONE) DIBLOCK COPOLYMER MICELLES: THEORETICAL AND EXPERIMENTAL DATA AND CORRELATIONS  Chapter 3 has been removed due to copyright restrictions. In this chapter five model hydrophobic drugs: indomethacin, curcumin, plumbagin, paclitaxel and etoposide were solubilized by a series of micelle forming, water soluble MePEG-b-PCL diblock copolymers with varying MePEG and PCL block lengths. Micellar encapsulation resulted in increased aqueous solubilities for all of the drugs investigated. Drug solubilization was related to the compatibility between the drug and PCL, as determined by the Flory- Huggins interaction parameter ("sp), as well as the PCL block length. This data was originally published in Letchford K et al., (2008) Solubilization of hydrophobic drugs by methoxy poly(ethylene glycol)-block-poly(caprolactone) diblock copolymer micelles: theoretical and experimental data and correlations. Journal of Pharmaceutical Sciences. 97(3) 1179-1190.  146  CHAPTER 4 IN VITRO HUMAN PLASMA DISTRIBUTION OF NANOPARTICULATE PACLITAXEL IS DEPENDENT ON THE PHYSICOCHEMICAL PROPERTIES OF METHOXY POLY(ETHYLENE GLYCOL)-BLOCK-POLY(!-CAPROLACTONE) NANOPARTICLES2  4.1 INTRODUCTION Studies in which amphiphilic copolymer nanoparticles loaded with drugs such as doxorubicin were administered intravenously, have demonstrated that these formulations produce increased circulation time and plasma half life, altered tissue distribution and in the case of tumour models, passive targeting via the enhanced permeation and retention effect (Alakhov et al., 1999; Kwon et al., 1994; Kwon et al., 1993; Yokoyama et al., 1991). Nanoparticulate formulations of the anticancer drug paclitaxel have also been investigated in order to improve its water solubility and toxicity profile. Paclitaxel is commercially formulated in a mixture of Cremophor" EL and ethanol, as Taxol#. However, Cremophor" EL has been shown to be responsible for hypersensitivity reactions (Dorr, 1994) and so there has been significant effort to find a replacement for this excipient. Copolymer micellar formulations readily solubilize PTX and are better tolerated than Taxol#, resulting in an increased maximum tolerated dose (Liggins and Burt, 2002). These formulations have demonstrated improved efficacy against several tumor types compared to Taxol# due to the ability to administer a higher dose of paclitaxel (Ahmed et al., 2006; Kim et al., 2001; Zhang et al., 1997). Some studies have  2  A version of this chapter has been submitted for publication. Letchford et al., (2008) In vitro human plasma distribution of nanoparticulate paclitaxel is dependent on the physicochemical properties of methoxy poly(ethylene glycol)-block-polycaprolactone nanoparticles. Eur, J. Pharm. Biopharm.  147 shown more favorable pharmacokinetics and biodistribution of micellar PTX compared to Taxol!, with longer circulation times (Hamaguchi et al., 2005; Han et al., 2006). On the other hand, in vivo evaluations of several micellar PTX formulations have shown little benefit over Taxol!, with respect to prolonging drug circulation time and increasing tumour uptake (Burt et al., 1999; Kim et al., 2001; Le Garrec et al., 2004).  Our group compared the in vitro distribution of PTX solubilized in MePEG-b-PDLLA micelles and free PTX into the component fractions of plasma. Both free and micellized PTX rapidly distributed into lipoprotein and lipoprotein deficient plasma fractions suggesting that the drug was released from the micelles and associated with the plasma proteins (Ramaswamy et al., 1997). Although pharmacokinetic studies of non-drug loaded diblock copolymer micelles have shown that copolymer nanoparticles are capable of prolonged circulation in the blood (Liu et al., 2007; Yamamoto Y, 2001), in many cases it is apparent that PTX does not remain with the carrier. It is evident that in order to increase the blood circulation time of PTX, better drug retention within the nanoparticle core is necessary. Several methods to retain PTX in the nanoparticle core have been investigated including cross-linking the hydrophobic blocks (Kim et al., 1999; Shuai et al., 2004), conjugating the drug to the core forming block (Zhang et al., 2005) and altering the structure of PTX so that it is more compatible with the micelle core (Forrest et al., 2008). It has also been suggested that increasing the hydrophobic block length may retain the drug for longer, by increasing thermodynamic and kinetic stability of the nanoparticle as well as prolonging the release of the drug (Allen et al., 1999).   148 Although it has been demonstrated that there are differences in the physicochemical properties of micelles and nanospheres, such as mobility of the core blocks and thermodynamic and chemical stabilities, there has not been a comparison of how these two types of nanoparticles differ in their interaction with blood and plasma constituents or how the encapsulated drug distributes into plasma fractions. Given the administration of drug loaded nanoparticles directly into the blood, it is also critical to understand the hemocompatibilities of these nanoparticles. In the work described in this chapter, two methoxy poly(ethylene glycol)-block-poly(caprolactone) (MePEG-b-PCL) amphiphilic diblock copolymers, one with a short PCL block (micelle-forming) and the other with a long PCL block (solid-like core, nanosphere-forming) were selected in order to elucidate the interactions of micelles and nanospheres with plasma components. Micelles and nanospheres were characterized with respect to their hydrodynamic diameters, core microviscosities, solubilization and in vitro release of PTX, and hemocompatibility, as assessed by measuring coagulation times and the rate of hemolysis. Relationships between these physicochemical characteristics and the differences observed in the in vitro plasma distribution of PTX solubilized by these two nanoparticulate delivery systems are described. 4.2 EXPERIMENTAL 4.2.1 Materials The synthesis and characterization of the two copolymers used in these studies, MePEG114-b-PCL19 and MePEG114-b-PCL104, were described in Chapter 2 (Table 2.1). 1,6-diphenyl-1,3,5-hexatriene (Sigma Aldrich Canada Ltd, Oakville, ON), sphingomyelin (Sigma Aldrich Canada Ltd) paclitaxel (Polymed Therapeutics Inc., Houston, TX), 3 H  149 paclitaxel (Moravek Biochemicals, Brea, CA) polysorbate 20 (Sigma Aldrich Canada Ltd) and Triton X 100 (Sigma Aldrich Canada Ltd) were used as supplied without further purification. The solvents, N,N-dimethyl formamide (Fisher Scientific Co., Ottawa, Ont.) acetonitrile (Fisher Scientific Co., HPLC grade), ethanol (Fisher Scientific Co., HPLC grade), methanol (Fisher Scientific Co., HPLC grade) and chloroform (Fisher Scientific Co., HPLC grade) were also used as supplied.  4.2.2 Formation and Characterization of MePEG-b-PCL Nanoparticles Nanoparticles were formed by a nanoprecipitation technique. Briefly, the copolymer was dissolved in N,N,-dimethyl formamide (DMF) and this solution was added drop wise to rapidly stirring PBS (0.01M, pH 7.4). The DMF was removed from the solution by dialysis in PBS overnight using 3500 MWCO Spectra/Por! dialysis membranes (Spectrum Laboratories, Inc., Rancho Dominguez, CA). The resulting dialysate was diluted with PBS so that the final copolymer concentration was 0.36 mM.  The hydrodynamic diameter of nanoparticles was determined by light scattering measurements carried out on a Malvern 3000HS Zetasizer (Malvern Instuments Ltd, Malvern, UK) with a 532 nm He-Ne laser and 90° collecting optics. Measurements were made at 37°C on copolymer samples at a concentrations ranging from 15 – 240 mg/ml of copolymer in the initial DMF solution. Data was analyzed using CONTIN algorithms provided by Malvern Zetasizer software.   150 Cryogenic transmission electron microscopy (cryo TEM) was carried out by Dr. Göran Karlsson at Uppsala University. Nanoparticles were prepared in distilled water at a copolymer concentration of 0.36 mM as previously described. Samples were equilibrated at 25°C and 99% relative humidity in a climate chamber prior to sample preparation. Copper grids (Agar Scientific, Essex, England) were covered with a perforated cellulose acetate butyrate film (Aldrich-Chemie, Steinheim, Germany) and a thin coating of carbon. A small amount of sample (<1 µl) was deposited on the grid followed by blotting with filter paper to leave a thin film on the grid. The grid was then vitrified in liquid ethane, held just above its freezing point of –182°C. A Zeiss EM 902A Transmission Electron Microscope (Carl Zeiss NTS, Oberkochen, Germany) was used to image the samples.  The instrument was operated at 80 kV and in zero loss bright-field mode. Digital images were recorded under low dose conditions with a BioVision Pro-SM Slow Scan CCD camera (Proscan GmbH, Scheuring, Germany) and iTEM software (Soft Imaging System, GmbH, Münster, Germany).  To investigate the differences in the microviscosity of the nanoparticle cores, the MePEG-b-PCL nanoparticles were loaded with the fluorescent probe 1,6-diphenyl-1,3,5- hexatriene (DPH) by adding 4.2 µL of a 2 µM solution in chloroform to the nanoparticle solution in amber glass vials. The probe was allowed to partition into the nanoparticle core overnight with stirring. For comparison, polysorbate 20 micelles were prepared by diluting polysorbate 20 in PBS so the final surfactant concentration was 6 mg/ml. Similar to the MeEPG-b-PCL nanoparticles, DPH was partitioned into polysorbate 20 micelles overnight with stirring. As a positive control, DPH was solubilized in a sphingomyelin  151 bilayer. Sphingomyelin in chloroform and the stock DPH solution were added to an amber glass vial and the solvent was evaporated under nitrogen gas to form a thin film. Warm PBS was added to the vial and the film was dissolved by vortexing and sonicating so the final sphinomyelin concentration was 6 mg/ml. For all samples, the fluorescent intensity of DPH was measured parallel (Ivv) and perpendicular (Ivh) to vertically polarized light over the temperature range of 27 – 61ºC using a Varian Cary Eclipse fluorescence spectrophotometer (Varian Inc., Palo Alto, CA). Excitation and emission intensities were 365 and 428 nm, respectively. The fluorescence anisotropy (r) was calculated as:  ! r = Ivv" Ivh Ivv+ 2Ivh   (4.1)  To determine the maximum loading of PTX in the nanoparticles, several nanoparticle solutions were prepared in which increasing amounts of drug were added to the DMF/copolymer solution prior to nanopreciptation. After dialysis, the dialysate was centrifuged at 14000 rpm to remove any precipitate and an aliquot of the nanoparticle solution was dried under a stream of nitrogen gas and reconstituted in acetonitrile. This solution was analyzed for PTX content by HPLC as previously described. The loading efficiency of the nanoparticles was calculated as:  ! Loading Efficiency = [PTX]so lub ilized [PTX]added "100  (4.2)   152 Where [PTX]solubilized is the concentration of PTX found in solution after encapsulation in nanoparticles as determined by HPLC and [PTX]added is the concentration of PTX added to solution during the formation of nanoparticles. 4.2.3 In Vitro Paclitaxel Release Nanoparticles with a copolymer concentration of 0.36 mM were loaded with PTX as described above, with the exception that a small amount of 3 H PTX was added to the drug/copolymer solution prior to dialysis. The initial cold PTX loading of nanoparticles was 40 µg/ml. Loaded nanoparticle solutions (3 ml) were added to 6000-8000 MWCO dialysis bags (Spectrum Laboratories, Inc., Rancho Dominguez, CA) and dialysed against 500 ml of PBS at 37ºC with shaking at 75 rpm. As a control, 3 ml of PBS in a dialysis bag was spiked with an aliquot of cold PTX in ethanol containing a trace amount of 3 H PTX so that the resulting drug concentration was 1 µg/ml, equivalent to the aqueous solubility of PTX. At time points of 0, 2, 4, 8, 12 hours and 1, 2, 4 and 7 days the volumes of the dialysis bags were measured and a 10 µl sample was taken for measurement of the remaining 3 H PTX in the dialysis bags. At each time point, the entire external release media was exchanged with fresh PBS. The concentration of 3 H PTX remaining in the dialysis bag at each time point was determined by beta scintillation counting (Beckman Coulter Canada, Mississagua, ON). The cumulative percent drug released was calculated by subtracting the amount of drug remaining from the initial amount of drug in the dialysis bag at the beginning of the experiment. The data were expressed as cumulative percentage of drug released as a function of time.   153 4.2.4 Blood Compatibility and In Vitro Paclitaxel Plasma Partitioning 4.2.4.1 Blood Coagulation Blood was collected from a healthy human volunteer in an evacuated siliconized glass tube containing 3.2% sodium citrate (protocols approved by UBC Clinical Research Ethics Board project # H07-02198 (Appendix 1)). Platelet poor plasma was isolated from the blood sample by centrifuging at 2000 g for 15 minutes at room temperature. A coagulation analyzer with mechanical endpoint determination (ST4 Diagnostica Stago, Inc, Parsippany, NJ) was used to analyze prothrombin time (PT) and activated partial thromboplastin time (APTT). Nanoparticle solutions with varying concentrations were added to the plasma samples in a 1:1 ratio and mixed at room temperature so that the final copolymer concentration varied from 0.0036 – 0.36 mM. For the PT determination, Innovin! reagent (Siemens Healthcare Diagnostics, Inc, Deerfield, IL) was added to the sample, whereas for APTT, partial thromboplastin reagent, Actin! (Siemens Healthcare Diagnostics, Inc, Deerfield, IL), was added along with calcium chloride. Controls consisted of identical volumes of PBS added to the plasma. Results were analyzed for statistical significance using an ANOVA. Differences were considered significant at p < 0.05. A Neuman-Keuls post hoc test was performed when a difference was detected.  4.2.4.2 Hemolysis Blood was collected from a healthy human volunteer in an evacuated siliconized glass tube containing sodium heparin. To separate red blood cells, the sample was centrifuged at 2000 rpm for 10 min and 1 ml of the packed cells was removed and washed 3 times with PBS. A stock suspension of erythrocytes in PBS was prepared so that the cell count  154 was 1x10 8  cells/ml. Nanoparticle dispersions at 0.02%, 0.1% and 0.2% w/v were prepared by nanoprecipitation as described above. As a positive control, micelles of a copolymer that has previously been shown to cause hemolysis, MePEG17-b-PCL4 (Zastre et al., 2007), were prepared by direct dissolution of the copolymer in 37ºC PBS at the concentrations mentioned above. An equal volume of erythrocyte suspension and nanoparticle solution was combined in a microcentrifuge tube so that the final erythrocyte concentration was 5x10 7  cells/ml and the final copolymer concentration was 0.1%, 0.05% or 0.01% w/v. A spontaneous hemolysis control was prepared by incubation of erythrocytes with PBS, also in a 1:1 ratio. A 100% hemolysis control was prepared by the addition of an equal volume of erythrocyte suspension and 1% Triton X-100. All tubes were incubated at 37ºC with tumbling. At time intervals of 1, 6 and 12 hours, the tubes were centrifuged at 20000 g for 15 seconds and the hemoglobin released in the supernatant was detected by UV absorbance at 540 nm (Cary Bio 50 UV Visible spectrophotometer, Varian). The percent hemolysis was calculated by:  ! %Hemolysis = (Abssample"Absspontaneous) Abs100%hemolysis #100  (4.3)  Where Abssample is the absorbance of the supernatant of the erythrocytes and nanoparticle suspension, Absspontaneous is the absorbance of the supernatant of the erythrocytes and PBS suspension and Abs100%hemolysis is the absorbance of the supernatant of the erythrocyte and Triton X-100 solution.   155 4.2.4.3 In Vitro Paclitaxel Plasma Distribution Nanoparticle solutions were prepared with a polymer concentration of 0.36 mM and a loading of 40 µg/ml of cold PTX along with a trace amount of 3H PTX as described above. An aliquot of these solutions or free PTX with a trace amount of 3 H PTX was added to human plasma (Bioreclamation Inc., Hicksville, NY) so that the final PTX concentration in plasma was 6.7 µg/ml. The samples were incubated for either 1, 6, 12 or 24 hours at 37°C with shaking at 75 rpm. After the specified incubation period, 1 ml of plasma was removed for comparison. Into the remaining 3 ml, 1.02 g of NaBr was dissolved to adjust the density of the plasma to 1.25 g/ml and the plasma was cooled at 4°C for 2 hours to prevent any further drug redistribution. The plasma was then separated into its lipoprotein and lipoprotein deficient fractions by density gradient ultracentrifugation as described elsewhere (Cassidy et al., 1998). Briefly, onto the plasma, sodium bromide solutions with densities of 1.21, 1.063 and 1.006 g/ml were layered, representing the high density lipoprotein (HDL) fraction, low density lipoprotein (LDL) fraction and very low density lipoprotein / chylomicron (VLDL) fractions, respectively. The solutions were centrifuged for 18 hrs at 40000 rpm (288000 g) in SW41 Ti swinging buckets at 15°C (Beckman Coulter, Mississauga, ON). Immediately after centrifugation, the layers were removed and the volumes measured. The amount of 3 H PTX in each fraction was quantified by beta scintillation counting and the amount of drug expressed as a percentage of total 3 H PTX as determined by the 1 ml of plasma removed prior to centrifugation. In order to determine whether the partitioning of the drug was due to interactions with plasma components and not due to the density of the formulation, plasma free controls were run. For these controls, the same experiment was run as  156 described above, except the drug loaded nanoparticles or free PTX were incubated for 1 hr in distilled water (in place of plasma). Results were analyzed for statistical significance using an ANOVA. Differences were considered significant at p < 0.05. A Neuman-Keuls post hoc test was performed when a difference was detected.  4.3 RESULTS 4.3.1 Formation and Characterization of MePEG-b-PCL Nanoparticles In order to eliminate the potential influence of different methods of preparation on the properties of the nanoparticles, the method of nanoprecipitation was used for both copolymers. The nanoparticulate dispersions formed by the two copolymers differed in appearance. Upon visual inspection, the MePEG114-b-PCL104 dispersion was opalescent in appearance, becoming increasingly turbid with concentration, while the MePEG114-b- PCL19 dispersion was clear and one phase over the range of copolymer concentrations tested. The diameters of the nanoparticles were determined by light scattering. Both types of nanoparticles had monodisperse distributions. The average hydrodynamic diameter of nanoparticles composed of MePEG114-b-PCL19 remained constant at approximately 40 nm over all concentrations tested, whereas those composed of MePEG114-b-PCL104 increased dramatically as the concentration of copolymer in the initial DMF solution increased, ranging from 73 – 135 nm over a copolymer concentration range of 15 – 180 mg/ml (Figure 4.1). Inspection of the cryo TEM images revealed spherical particles; however, the diameters of the particles were smaller than those determined by DLS, particularly for the MePEG114-b-PCL19 micelles (Figure 4.2).   157 Fluorescence anisotropy was used to investigate the differences in the core microviscosity of nanoparticles formed from either MePEG114-b-PCL104 or MePEG114-b- PCL19 (Figure 4.3). The hydrophobic fluorescence probe, DPH, is sensitive to the microviscosity of its surroundings and displays changes in its depolarization as measured by fluorescence anisotropy. As a positive control, DPH was solubilized in a sphingomyelin bilayer. Upon heating the sphinomyelin, the anisotropy dropped significantly between 35 and 45ºC. Both copolymer nanoparticles displayed a decrease in the anisotropy implying a decrease in the microviscosity of the core during heating. However, nanoparticles of MePEG114-b-PCL104 had higher anisotropy values than those composed of MePEG114-b-PCL19 over all temperature ranges. As a comparison, DPH was solubilized in polysorbate 20 micelles. Anisotropy values lower than those of MePEG114- b-PCL19 were found at all temperatures for polysorbate 20.  158                      Figure 4.1 Dependence of nanoparticle hydrodynamic diameter on copolymer concentration for MePEG114-b-PCL19 micelles (!) and MePEG114-b- PCL104 nanospheres ("). All points represent the mean hydrodynamic diameter ± SD (n = 3).  159                                           Figure 4.2 Cryo TEM images of (A) MePEG114-b-PCL19 micelles and (B) MePEG114- b-PCL104 nanospheres. Bar represents 100 nm. B A  160                     Figure 4.3 Anisotropy as a function of temperature. Spingomyelin bilayer (!) MePEG114-b-PCL19 micelles (") MePEG114-b-PCL104 nanospheres (#) polysorbate 20 micelles ($). All points represent the mean anisotropy ± SD (n = 3).  161 The solubilization capacity of PTX in nanoparticulate dispersions of each of the synthesized copolymers was investigated (Figure 4.4). In these experiments, each copolymer at a concentration of 0.36 mM was used to solubilize increasing amounts of drug. All solutions were clear up to a threshold amount of total drug added, after which precipitation was observed upon visual inspection. The amount of drug solubilized increased up to a maximum concentration, above which the solubilized concentration decreased significantly corresponding with the formation of a precipitate. The MePEG114- b-PCL19 micelles solubilized PTX at a concentration of 40 µg/ml (1.6% w/w) at approximately an 80% loading efficiency, whereas MePEG114-b-PCL104 nanospheres solubilized the drug at 180 µg/ml (3% w/w) with a 90% loading efficiency. 4.3.2 In Vitro Paclitaxel Release The in vitro release of PTX from nanoparticles at a copolymer concentration of 0.36 mM and a drug loading of 40 µg/ml for both copolymers was investigated (figure 4.5). For both types of nanoparticles, PTX was released in a controlled and sustained fashion, with MePEG114-b-PCL19 micelles releasing 92% of their drug payload after 7 days whereas approximately 60% was released from MePEG114-b-PCL104 nanospheres. Free PTX was released rapidly from the dialysis bags with complete release of the drug by 12 hours.  162      Figure 4.4 Solubilization of PTX by (A) MePEG114-b-PCL19 micelles and (B) MePEG114-b-PCL104 nanospheres formed by nanoprecipitation of copolymer and drug in DMF solutions. All nanoparticles were formed in PBS (0.01M, pH 7.4) using copolymer concentrations of 0.36mM. [PTX] solubilized (!), loading efficiency (!). Data points and error bars represent the mean ± SD (n=3). A B  163                     Figure 4.5 Release of PTX from dialysis bags. Drug was either in its free form (!) or solubilized in MePEG114-b-PCL19 micelles (") or MePEG114-b-PCL104 nanospheres (#) and released into PBS (0.01M, pH 7.4) at 37°C. Data points and error bars represent the mean ± SD (n=3).  164  4.3.3 Blood Compatibility and In Vitro Paclitaxel Plasma Distribution 4.3.3.1 Blood Coagulation Coagulation of human plasma was assessed in the presence of MePEG114-b-PCL19 micelles or MePEG114-b-PCL104 nanospheres with varying copolymer concentrations (Figure 4.6). Changes in the extrinsic and common coagulation pathways were evaluated by measurements of the prothrombin time (PT). No statistical difference in PT was found for plasma incubated with either type of nanoparticle as compared to the PBS control. The intrinsic pathway was evaluated using activated partial thromboplastin time (APTT). The APTT was decreased for MePEG114-b-PCL19 at the highest concentration tested, 0.36mM, and increased for MePEG114-b-PCL104 at 0.0036mM, as compared to the PBS control. 4.3.3.2 Hemolysis Neither MePEG114-b-PCL19 micelles nor MePEG114-b-PCL104 nanospheres induced significant amounts of hemolysis of erythrocytes over a 12 hour incubation time at concentrations below 0.1% w/v. The positive control, MePEG17-b-PCL4, resulted in approximately 30% hemolysis at a concentration of 0.05% w/v. When the polymer concentration was increased to 0.1% w/v there was a small amount of hemolysis (5%) by MePEG114-b-PCL104 nanospheres at 12 hours. However this level of hemolysis was small compared to the considerable amount of hemolysis (40% at 6 hours) that was observed for micellar solutions of MePEG17-b-PCL4 at this concentration (Figure 4.7).  165    Figure 4.6 Effect of nanoparticle type and concentration on the (A) prothrombin time (PT) and (B) activated partial thromboplastin time (APTT). White bars represents control (PBS), black bars represents MePEG114-b-PCL19 micelles, grey bars represents MePEG114-b-PCL104 nanospheres. All bars represent mean clotting time ± SD (n=3). * p <0.05, different from control. A B  166                        Figure 4.7 Time course of hemolysis by MePEG17-b-PCL4 micelles (!), MePEG114-b- PCL19 micelles ("), and MePEG114-b-PCL104 nanospheres (#) at 37°C at copolymer concentrations of 0.1% w/v. Data represents mean ± SD (n = 3).  167  4.3.3.3 In Vitro PTX Plasma Distribution The distribution of PTX in plasma as a function of its formulation was examined. The control groups consisted of PTX loaded micelles, nanospheres or free PTX incubated in distilled water. All fractions in the control groups were clear and colourless with the exception of the samples containing the MePEG114-b-PCL104 nanospheres. In these samples, an opalescent layer was present in the 1.063 g/ml fraction, close to the interface with the 1.21 g/ml fraction. This layer was termed the nanoparticle fraction (“NP fraction”) and its position was marked on the outside of the tube and carefully separated from the remainder of the 1.063 g/ml fraction and analyzed for 3 H PTX (see schematic inset in Figure 4.8). Although the “NP fraction” was not visually present in the micellar and free PTX samples, in order to remain consistent, the equivalent layer was removed from these samples and analyzed for 3 H PTX. It was found that 100% of the 3 H PTX formulated in MePEG114-b-PCL104 and 76% of the 3 H PTX formulated in MePEG114-b- PCL19 was recovered in the “NP fraction” in distilled water (Figure 4.8).  Following density gradient ultracentrifugation of free PTX or micellar PTX incubated in plasma, four visibly distinct layers were observed representing, from top to bottom, the VLDL, LDL, HDL and LPDP fractions, corresponding to the 1.006 g/ml, 1.063 g/ml, 1.21 g/ml and 1.25 g/ml density layers, respectively. After centrifugation of the MePEG114-b-PCL104 samples, again, the thin, opalescent “NP fraction” at the bottom of the 1.063 g/ml (LDL) fraction, close to the interface with 1.21 g/ml (HDL) density layer was observed. This layer was carefully separated and analyzed for 3 H PTX content along  168 with the other layers. Again, to remain consistent, this same fraction was also removed from the plasma samples containing free PTX and MePEG114-b-PCL19 micelles and termed the “NP fraction”, even though this layer was clear in these samples. After incubation of free PTX in human plasma, approximately 50% of the 3 H PTX was found in the LPDP fraction and the remaining drug was equally distributed among the other fractions with the exception of the VLDL fraction, which contained very little drug (Figure 4.9). The plasma distribution of PTX solubilized in MePEG114-b-PCL19 micelles was similar to that of free PTX, with approximately 50% of the drug found in the LPDP fraction and an equal distribution of the remaining drug in the other fractions (Figure 4.10). For PTX formulated in MePEG114-b-PCL104 nanospheres, approximately 45% of the radiolabelled drug was found in the “NP fraction” with approximately 35% of the remaining 3 H PTX found in the LPDP fraction (Figure 4.11). To further investigate this effect, free drug was first incubated in plasma for 10 minutes followed by addition of blank MePEG114-b-PCL104 nanospheres and an additional 1 hour incubation. The 3 H PTX distribution was similar to the data for the previous experiment in which most of the 3 H PTX was found in the “NP fraction” and the majority of the remaining drug in the LPDP fraction (Figure 4.11). There were some statistically significant changes in the distribution of free and nanoparticulate PTX over the incubation times. In the absence of any clear trend in the data, these differences were not given further consideration.  169              Figure 4.8 Partitioning of PTX into various density fractions after a 1 hr incubation in distilled water. Drug was formulated in either MePEG114-b-PCL19 micelles (white) or MePEG114-b-PCL104 nanospheres (black). All values represent mean ± SD (n=6). Inset is a schematic of an ultracentrifuge tube with the position of the various density fractions and their corresponding lipoprotein fractions after ultracentrifugation. NP fraction represents the clearly separated white band of MePEG114-b-PCL104 nanospheres.  170                  Figure 4.9 Time dependence of partitioning of free PTX into various human plasma fractions during incubation times of 1 hr (black), 6 hrs (white), 12 hrs (grey), and 24hrs (diagonal stripes). All values represent mean ± SD (n=6). * p < 0.05, different from all other samples in plasma fraction.  171                Figure 4.10 Time dependence of partitioning of MePEG114-b-PCL19 micellar PTX into various human plasma fractions during incubation times of 1 hr (black), 6 hrs (white), 12 hrs (grey), and 24hrs (diagonal stripes). All values represent mean ± SD (n=6) All values represent mean ± SD (n=6). # p < 0.05, different from all other samples in plasma fraction. * p < 0.05, different from 1 hr sample in plasma fraction.  172              Figure 4.11 Time dependence of partitioning of MePEG114-b-PCL104 nanosphere PTX into various human plasma fractions during incubation times of 1 hr (black), 6 hrs (white), 12 hrs (grey), and 24hrs (diagonal stripes). Hatched bars represent distribution of PTX after incubation of free PTX for 10 min followed by addition of blank MePEG114-b-PCL104 nanospheres and an additional 1 hr incubation. All values represent mean ± SD (n=6). * p < 0.05, different from all other samples in plasma fraction. # p < 0.05, different from 1 hr sample in plasma fraction.  173  4.4 DISCUSSION Nanoparticles were formed by the solvent displacement method, also referred to as nanoprecipitation. This method was chosen due to the low aqueous solubility of the MePEG114-b-PCL104 copolymer requiring that the copolymer first be dissolved in an organic solvent. Although MePEG114-b-PCL19 was readily water-soluble, which allows for the formation of micelles by direct dissolution, to maintain consistent preparation methods for both copolymers, MePEG114-b-PCL19 micelles were formed by nanoprecipitation. The hydrodynamic diameters of the formed nanoparticles at a concentration of 0.36 mM, as determined by DLS, were found to be dependent on the length of the hydrophobic block. Nanospheres composed of MePEG114-b-PCL104 had a diameter of 80 nm, approximately twice as large as the MePEG114-b-PCL19 micelles. Electron micrographs revealed that the nanoparticles had a spherical morphology; however, the diameters of the nanoparticles did not agree with the light scattering results. By TEM, the nanoparticles appeared to be significantly smaller, particularly for the MePEG114-b-PCL19 micelles. This observation has been described by other research groups and was rationalized by the fact that only the dense core of the nanoparticles is evident in the micrographs (Hans et al., 2005). Since the core block of the MePEG114-b- PCL19 micelles is relatively short, it is to be expected that the diameter of the micelles in the TEM image is considerably smaller than that determined by DLS. Interestingly, the hydrodynamic diameter of the nanospheres formed from MePEG114-b-PCL104 had a clear dependence on the concentration of copolymer used in their formation. Riley et al. also observed a concentration dependence of the hydrodynamic diameter of MePEG-b-  174 PDLLA nanoparticles with long PDLLA blocks (Riley et al., 2001). In these studies, a series of MePEG-b-PDLLA copolymers with a fixed MePEG block length and increasing PDLLA molecular weights were synthesized. It was observed that the nanoparticles formed by copolymers with relatively low PDLLA molecular weights had hydrodynamic radii dependent only on the hydrophobic block length and this radius scaled linearly with the number of hydrophobic blocks, characteristic of block copolymer micelles. Conversely, once the hydrophobic block reached a critical length, the hydrodynamic radii of the nanoparticles became highly dependent on the concentration of copolymer used in the organic phase prior to nanoprecipitation, similar to nanoparticles formed from homopolymers.  Using fluorescence anisotropy, we showed that as the PCL block length increased there was a significant increase in the microviscosity of the nanoparticle cores. It has been documented that the fluorescence probe DPH partitions into the hydrophobic domain of membranes and core region of micelles (Zastre et al., 2002; Zhang et al., 1996). This probe is sensitive to the microviscosity of its surroundings and displays changes in its depolarization as measured by fluorescence anisotropy. A smaller value of anisotropy indicates a lower degree of depolarization and therefore a less viscous environment. As a positive control, DPH was solubilized in a sphingomyelin bilayer. Upon heating the sphingomyelin, the anisotropy value dropped significantly between 33°C and 47ºC, with a midpoint at approximately 40°C, characteristic of the phase transition of the bilayer from a gel to a less viscous liquid crystalline state (Kuikka et al., 2001). Both copolymer nanoparticles displayed a slight decrease in the anisotropy implying a decrease in the  175 microviscosity of the core during heating. However, MePEG114-b-PCL104 nanospheres were more viscous than MePEG114-b-PCL19 micelles over the whole temperature range. As a comparison, DPH was solubilized in polysorbate 20 micelles. Anisotropy values for MePEG114-b-PCL19 were considerably higher than those of the short chain surfactant micelles of polysorbate 20, indicating a higher viscosity core, presumably due to the longer, entangled chains of the PCL. Similar to our findings, Lavasanifar et al. used the hydrophobic, fluorescent probe 1,3-(1,1’-dipyrenyl)propane to show that increased poly(L-amino acid) (PLAA) block lengths induced a noticeable increase in the core microviscosity of MePEG-b-PLAA micelles (Lavasanifar et al., 2001). Studies by Heald et al. used 1 H NMR techniques to show that as the PDLLA block of MePEG-b-PDLLA copolymers increased, the hydrophobic core became more solid-like (Heald et al., 2002). Thus, our findings confirm those of other researchers, suggesting that the microviscosity of the hydrophobic core is correlated to the length of the core-forming blocks. The short hydrophobic block, water-soluble MePEG114-b-PCL19 copolymer, formed micelles with a relatively low viscosity core. Conversely, the more hydrophobic MePEG114-b-PCL104 copolymer formed nanospheres with a more solid-like core characterized by a higher microviscosity.  The PTX solubilization capacity of nanoparticulate dispersions of each of the synthesized copolymers was investigated. For each solution, the amount of drug solubilized increased up to a maximum concentration, above which drug precipitation occurred with a corresponding decrease in the solubilized drug concentration and loading efficiency (Figure 4.4). The MePEG114-b-PCL104 nanosphere formulation solubilized more drug at a  176 higher loading efficiency than that of the MePEG114-b-PCL19 micelles, which may be attributed to the larger core volume of MePEG114-b-PCL104 nanospheres as indicated by the TEM micrographs and hydrodynamic diameters. A review of the literature indicates that the amount of PTX solubilized by diblock copolymers can range from 0.5% up to 25% w/w (Lee et al., 2003; Zhang et al., 1996). A number of factors may explain this broad range of PTX loadings, including differences in hydrophobic block length and preparation method. The compatibility between the drug and core-forming block has also been shown to be a contributing factor in the amount of drug that can be solubilized by nanoparticles. Lui et al. demonstrated that the hydrophobic anticancer compound, ellipticine, was solubilized by MePEG-b-PCL to a greater extent than by MePEG-b- PDLLA (Liu et al., 2004). This group attributed the difference to a greater degree of compatibility between PCL and the drug as calculated by solubility parameters and the enthalpy of mixing. In Chapter 3, it was demonstrated that the amount of drug solubilized by MePEG-b-PCL micelles was directly correlated to the compatibility between the hydrophobic block and drug, as determined by the calculated Flory-Huggins interaction parameter. Furthermore, our compatibility calculations predicted that PTX was not highly compatible with PCL, which was demonstrated experimentally by a low solubilization of PTX. Zhang et al. were able to achieve PTX loading levels of 25% w/w in MePEG-b- PDLLA micelles using the film hydration method (Zhang et al., 1996). In the current study, we were not able to attain as high of a drug loading, likely due a lower compatibility between PTX and PCL as compared to PTX and PDLLA. In addition, the method of preparation can affect drug loading. In Chapter 3, we were able to solubilize more PTX in MePEG-b-PCL micelles using a film hydration method. However, film  177 hydration of water insoluble copolymers is not possible, therefore, for the current study, dialysis was necessary to produce nanospheres from the MePEG114-b-PCL104 copolymer.  Free PTX was released from the dialysis bags at a rapid rate with complete depletion of the drug from the bags by 12 hours, thus demonstrating that the dialysis membrane would not retard the release of the drug from nanoparticles to any great extent. Paclitaxel was released from both formulations in a controlled and sustained manner over greater than 7 days. The rate of release of PTX was faster from MePEG114-b-PCL19 micelles, with nearly all of the drug released by 7 days, compared to the MePEG114-b-PCL104 nanospheres, which had released approximately 60% of loaded drug in 7 days. Other groups have also found that an increase in the hydrophobic block decreases the release rate of the encapsulated drug. Deng and coworkers found that a increase in the molecular weight of the PLLA block of MePEG-PLLA diblock copolymers from 3515 g/mol to 4247 g/mol resulted in an decrease in the cumulative PTX release of almost 10% over 4 days (Deng et al., 2005). Likewise, Kim et al. showed that increases in the PCL block length decreased the release rate of indomethacin (Kim et al., 1998). In both of these cases, the decrease in the release rate was attributed to the increased interaction between the longer hydrophobic block of the nanoparticles and the drug. The drug release from MePEG114-b-PCL104 nanospheres was characterized by a large amount of variability compared to that of the micelles. It is suggested that this may be due to the larger polydispersity index of the MePEG114-b-PCL104 copolymer (Table 2.1) resulting in a range of nanosphere diameters and release rates.   178 It is evident that a systemically administered (intravenous route) drug carrier, the system must be hemocompatible. For a range of concentrations, the two copolymers were tested for compatibility with human blood by investigating their ability to alter coagulation times and cause hemolysis.  The effects on coagulation were evaluated by analyzing the activated partial thromboplastin time (APTT) and the prothrombin time (PT). The APTT was found to be significantly decreased for micelles at the highest concentration and increased at the lowest nanospheres concentration as compared to the PBS control. These differences were considered to be of minor concern as the experimental values were well within the normal range of coagulation times of 32-46 seconds for APTT (Perkins, 2004). The lack of activation of the coagulation system may be attributed to the highly hydrated PEG brush-like covering of the surfaces of nanoparticles as PEGylated nanoparticle surfaces have been shown to inhibit adsorption of various plasma proteins, including plasma factors of the coagulation system, thus, preventing the activation of the coagulation cascade (Gref et al., 2000; Sahli et al., 1997).  Of critical importance for systemically administered amphiphilic copolymer systems is their hemolytic potential. It is well known that surfactants at sufficiently high concentrations are capable of disrupting cell membranes, such as those of erythrocytes, by penetration and saturation of the membrane with unimers followed by solubilization of the membrane lipids and proteins (Jones, 1999; Shalel et al., 2003). As shown previously by our group, the short block length amphiphilic copolymer MePEG17-b-PCL4 was capable of inducing rapid and significant hemolysis (Zastre et al., 2007). In the current study, there was no significant hemolysis when erythrocytes were incubated with either  179 the MePEG114-b-PCL19 micelles or MePEG114-b-PCL104 nanospheres. In Chapter 2, it was shown that MePEG114-b-PCL19 has a relatively low critical micelle concentration of 6.29 x 10 -7  M indicating a low number of unimers in solution and a high degree of thermodynamic stability (Letchford et al., 2008). We suggest that the stable core and well-hydrated corona of these nanoparticles would result in intact nanoparticles with few available unimers available for penetration into the cell membrane, thus, reducing membrane perturbation and subsequent hemolysis.  Free paclitaxel is highly bound to plasma proteins, specifically albumin and !1-acid glycoprotein (Kumar et al., 1993). In this study, the incubation of free PTX led to the partitioning of the drug equally into the LPDP (i.e. primarily albumin and !-1- glycoprotein) and lipoprotein fractions. The distribution of the free drug did not change over the incubation times tested, indicative of a rapid distribution of PTX among the plasma components, and confirming our earlier findings (Ramaswamy et al., 1997). When PTX encapsulated in MePEG114-b-PCL19 micelles was incubated in distilled water and separated by density gradient centrifugation, the majority of PTX was found in the lower portion of the 1.063 g/ml density fraction, or what we termed the “NP fraction” (Figure 4.8). We believe that the PTX found in this density fraction primarily remained associated with the micelles since the final concentration of copolymer after incubation, layering and centrifugation was still well above the CMC previously determined for this copolymer (Chapter 2, Table 2.1). Incubation of the micelles in plasma dramatically changed the distribution of the drug so that 50% of the drug was associated with the LPDP fraction, similar to that of free PTX. Since the copolymer concentration was  180 maintained well above the CMC, we believe this effect was not caused by destabilization of the micelles due to dilution, but rather from an interaction with plasma components. A similar drug distribution was found previously when PTX was solubilized by MePEG-b- PDLLA micelles (Ramaswamy et al., 1997). These findings suggest a few possibilities. Upon incubation in plasma, it is possible that there was disruption of the micellar structure and/or rapid drug release and partitioning into the LPDP and lipoprotein fractions. The other possibility is that the intact micelles with their drug payload were associated with the plasma fractions. Due to the fact that we were not able to assay the copolymer in the various plasma fractions, it was not possible to determine whether the drug was still associated with the copolymer. Previously, our group investigated the biodistribution of 3 H PTX solubilized in 14 C MePEG-b-PDLLA micelles in male Sprague-Dawley rats (Burt et al., 1999). It was found that PTX rapidly dissociated from the micellar components and the copolymer accumulated in the kidneys and bladder resulting in copolymer elimination through the urine. Conversely, the PTX was widely distributed among the tissues and the majority of the drug was eliminated in the feces. These results provide strong evidence for the likelihood of rapid release of the micellar drug during incubation in plasma, resulting in similar plasma profiles for free and micellar PTX.  When PTX encapsulated in MePEG114-b-PCL104 nanospheres was incubated in distilled water, it was possible to visualize and separate out the nanospheres from the other density fractions since the nanospheres centrifuged into a separate, concentrated layer (“NP fraction”) located between the 1.063 g/ml and 1.21 g/ml density fractions. It was found  181 that 100% of the encapsulated 3 H PTX was recovered from this fraction demonstrating that the encapsulated drug remained associated with the MePEG114-b-PCL104 copolymer during the ultracentrifugation process. Upon incubation in plasma, again the nanospheres were found in the same density fraction as when they were incubated in distilled water. The largest amount of the 3 H PTX was found in this layer with the majority of the remaining drug in the LPDP fraction. Since the nanoparticles were separated due to their density and this separation did not change when incubated in plasma, it was not possible to establish whether the nanospheres were associated with lipoproteins in the plasma prior to ultracentrifugation. However, the key finding is that the MePEG114-b-PCL104 nanospheres retained more of their drug payload, releasing less of it to the LPDP fraction, compared to MePEG114-b-PCL19 micelles or free drug. Interestingly, even when free PTX was incubated in plasma, followed by the addition of blank (non drug loaded) MePEG114- b-PCL104 nanospheres, again the majority of the drug was found in the NP fraction, indicating a high affinity of PTX for association with these nanospheres.  4.5 CONCLUSIONS In this study we have shown that increasing the hydrophobic block length of MePEG-b- PCL copolymers results in a dramatic shift in the physicochemical properties of nanoparticles made from this polymer. The readily water-soluble MePEG114-b-PCL19 formed micellar type nanoparticles with a small average hydrodynamic diameter that did not vary with copolymer concentration and possessed a low core microviscosity. The water insoluble MePEG114-b-PCL104 formed what we termed nanospheres, characterized by a larger hydrodynamic diameter that increased with copolymer concentration and a  182 more solid-like core. Compared to the micelles, nanospheres solubilized more PTX and released it at a slower rate. There was no major difference in the hemocompatibility between the two types of nanoparticles as determined by coagulation times and hemolysis. When PTX loaded micelles and nanospheres were incubated in human plasma, the majority of the micellar drug was associated with the lipoprotein deficient plasma fraction (LPDP), similar to the plasma distribution profile of free PTX. On the other hand, the majority of PTX solubilized in nanospheres, was found associated with these nanoparticles in a distinct density fraction, well separated from the LPDP fraction. These results demonstrate that the nanospheres have a high affinity for PTX allowing them to retain their drug payload when incubated in human plasma as compared to micelles.  183  4.6 REFERENCES  Ahmed F, Pakunlu R I, Brannan A, Bates F, Minko T and Discher D E. Biodegradable polymersomes loaded with both paclitaxel and doxorubicin permeate and shrink tumors, inducing apoptosis in proportion to accumulated drug. 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Physicochemical Evaluation of Nanoparticles Assembled from Poly(lactic acid)-Poly(ethylene glycol) (PLA-PEG) Block Copolymers as Drug Delivery Vehicles. Langmuir 17 (2001) 3168-3174. Sahli H, Tapon-Bretaudiere J, Fischer A-M, Sternberg C, Spenlehauer G, Verrecchia T and Labarre D. Interactions of poly(lactic acid) and poly(lactic acid-co-ethylene oxide) nanoparticles with the plasma factors of the coagulation system. Biomaterials 18 (1997) 281-288. Shalel S, Streichman S and Marmur A. Modeling surfactant-induced hemolysis by Weibull survival analysis. Colloids and Surfaces, B: Biointerfaces 27 (2003) 223-229. Shuai X, Merdan T, Schaper A K, Xi F and Kissel T. Core-Cross-Linked Polymeric Micelles as Paclitaxel Carriers. Bioconjuagte Chemistry 15 (2004) 441-448. Yamamoto Y N Y, Kato Y, Sugiyama Y, Kataoka K. Long-circulating poly(ethylene glycol)-poly(D,L-lactide) Block coolymer micelles with modulated surface charge. Journal of Controlled Release 77 (2001) 27-38. Yokoyama M, Okano T, Sakurai Y, Ekimoto H, Shibazaki C and Kataoka K. Toxicity and antitumor activity against solid tumors of micelle-forming polymeric anticancer drug and its extremely long circulation in blood. Cancer Research 51 (1991) 3229-3236. Zastre J, Jackson J, Bajwa M, Liggins R, Iqbal F and Burt H. Enhanced cellular accumulation of a P-glycoprotein substrate, rhodamine-123, by caco-2 cells using low molecular weight methoxypolyethylene glycol-block-polycaprolactone diblock copolymers. European Journal of Pharmaceutics and Biopharmaceutics 54 (2002) 299- 309. Zastre J, Jackson J K, Wong W and Burt H M. Methoxypolyethylene glycol-block- polycaprolactone diblock copolymers reduce P-glycoprotein efflux in the absence of a membrane fluidization effect while stimulating P-glycoprotein ATPase activity. Journal of Pharmaceutical Sciences 96 (2007) 864-875. Zhang X, Burt H M, Mangold G, Dexter D, Von Hoff D, Mayer L and Hunter W L. Anti- tumor efficacy and biodistribution of intravenous polymeric micellar paclitaxel. Anticancer Drugs 8 (1997) 696-701.  187 Zhang X, Jackson J K and Burt H M. Determination of surfactant critical micelle concentration by a novel fluorescence depolarization technique. Journal of Biochemical and Biophysical Methods 31 (1996) 145-150. Zhang X, Jackson J K and Burt H M. Development of amphiphilic diblock copolymers as micellar carriers of taxol. International Journal of Pharmaceutics 132 (1996) 195-206. Zhang X, Li Y, Chen X, Wang X, Xu X, Liang Q, Hu J and Jing X. Synthesis and characterization of the paclitaxel/MPEG-PLA block copolymer conjugate. Biomaterials 26 (2005) 2121-2128.  188  CHAPTER 5 SUMMARIZING DISCUSSION, CONCLUSIONS AND SUGGESTIONS FOR FUTURE WORK 5.1 SUMMARIZING DISCUSSION AND CONCLUSIONS Amphiphilic block copolymer nanoparticles have become a class of drug delivery systems of significant research interest. Due to the unique structure of these copolymers, they are capable of self assembling to form nanoparticles characterized by a hydrophobic core, surrounded by a highly water bound corona (Jones and Leroux, 1999). The core of these nanoparticles has been demonstrated to effectively solubilize a number of hydrophobic drugs providing an effective method for increasing the aqueous solubility of these compounds (Kwon, 1998). The corona acts to prevent the opsonization and subsequent recognition of the particles by the reticuloendothelial system, thus, prolonging their circulation in blood (Allen et al., 1999). Furthermore, due to their nanoscopic size, they can passively target disease sites that are characterized by leaky vasculature, such as tumors, via the enhanced permeation and retention effect (Kwon et al., 1994).  Due to these benefits, amphiphilic block copolymer nanoparticles are actively being investigated for the delivery of a number of hydrophobic anticancer drugs, including paclitaxel (PTX) (Deng et al., 2005; Kim et al., 2001). Due to some toxicities demonstrated by Cremophor! EL (Dorr, 1994), the excipient currently used to solubilize PTX, there has been considerable interest in the development of safer solubilization strategies. Biocompatible amphiphilic block copolymer micelles have been extremely successful in solubilizing PTX (Zhang et al., 1996); however, as suggested by in vitro  189 plasma distribution studies (Ramaswamy et al., 1997) and in vivo pharmacokinetic and biodistribution studies, the drug often does not remain encapsulated within the carrier and hence, prolonged circulation and passive targeting are not attained (Burt et al., 1999; Le Garrec et al., 2004). Therefore, it was the overall goal of this project to develop a non- toxic, amphiphilic block copolymer nanoparticulate delivery system that would not only increase the aqueous solubility of PTX, but also retain the drug in plasma. Such a formulation should provide prolonged circulation to enable the passive targeting of the drug delivery system to tumor sites.  It has been shown that the physicochemical properties of amphiphilic block copolymers, as well as the nanoparticulate delivery systems they form, are highly dependent on the relative block lengths of the copolymer (Heald et al., 2002; Riley et al., 2001). Therefore, to develop an amphiphilic block copolymer nanoparticulate drug delivery system that will effectively solubilize and retain PTX, it is essential to understand how the composition of the block copolymer relates to the performance of the nanoparticles. In order to do so, a series of amphiphilic block copolymers with varying hydrophilic and hydrophobic block lengths were synthesized. The physicochemical properties including, thermodynamic stability, phase behavior, particle size and drug solubilization were assessed to aid in the selection of copolymers for further evaluation of hemocompatibility and in vitro transport and partitioning of PTX in plasma components.  MePEG-b-PCL diblock copolymers were selected for formulating the nanoparticulate drug delivery system. MePEG-b-PCL has been used for the solubilization of several  190 hydrophobic drugs including rapamycin (Forrest et al., 2006), indomethacin (Kim et al., 2001), and cyclosporine A (Aliabadi et al., 2005). We considered this copolymer to be an appropriate candidate for drug delivery applications as PCL is biocompatible and has a history of use in biomedical applications including implants and sutures (Nair and Laurencin, 2007). Its high degree of hydrophobicity and prolonged degradation time make PCL a good choice for controlled drug delivery. Furthermore, our group has a particular interest in drug formulations based on MePEG-b-PCL amphiphilic diblock copolymers, based on an earlier observation that some members of a series of short block length MePEG-b-PCL copolymers reduced the efflux transport of some hydrophobic drug substrates by P-glycoprotein (Pgp), a protein that is implicated in mechanisms of drug resistance in tumor cells and decreased drug absorption through Pgp overexpressing epithelia (Zastre et al., 2002).   The synthesis of a series of MePEG-b-PCL diblock copolymers with varying hydrophilic and hydrophobic block lengths was reported in Chapter 2. This series formed a library of copolymers from which members were selected to investigate the effects of changes in hydrophilic and hydrophobic block lengths on the physicochemical properties of the copolymers and the micellar systems that are formed. The synthesis of the copolymers involved a ring opening reaction of !-caprolactone using the hydroxyl terminus of MePEG as the initiator. The reaction proved to be reliable with block lengths and molecular weights of the products predicted by the initial feed ratio of the reactants. Side products or residual reactants were undetectable by GPC and NMR. The resulting copolymers had MePEG blocks of 17, 44 or 114 repeat units and PCL blocks ranging  191 from 2 to 104 repeat units. All the synthesized copolymers were water soluble, with the exception of MePEG114-b-PCL104, which was extremely insoluble due to the large proportion of PCL.  Using a steady state pyrene fluorescence method, the CMC and Kv values were determined for the copolymers. These values provide an estimate of the thermodynamic stability of the micelles upon dilution as well as the degree of hydrophobicity of the core, respectively. An increase in the PCL block length resulted in lower CMCs and increased Kv values, due to the hydrophobic nature of PCL. The hydrodynamic diameter of the micelles, determined by light scattering, was also found to increase with corresponding increases in the PCL block length. This trend was obvious within copolymer series with the same MePEG molecular weight and was attributed to increases in the micelle core volume.  The series of copolymers based on MePEG with a molecular weight of 750 g/mol displayed temperature dependent phase behavior, not normally found for amphiphilic copolymers. Characterization of this behavior revealed that these copolymers formed clear micellar solutions between two temperature boundaries, the lower being the Krafft point and the upper, the cloud point. Dependence of these phase transitions on the PCL block length was found to be due to the decreased aqueous solubility of the copolymers as the PCL block length increased. Copolymers with shorter PCL blocks had lower Krafft points and higher cloud points compared to the longer PCL block copolymers.   192 It is important to be able to efficiently solubilize high concentrations of hydrophobic drugs and retain the payload within nanoparticles to enable prolonged blood circulation and passive targeting. The compatibility between the drug and the core-forming block is a key factor determining the amount of drug solubilized and its release rate (Allen et al., 1999). In Chapter 3, the solubilization of several hydrophobic drugs by the series of water soluble MePEG-b-PCL copolymers was investigated. Empirical relationships between the amount of drug solubilized and the compatibility between the drug and core-forming block, as determined by the Flory-Huggins interaction parameter (!sp), were developed. Furthermore, the effect of changes in the core-forming block length on drug solubilization were also investigated. The compatibility of the drug with the core-forming block is often qualitatively used as an explanation for the increased solubilization of one drug over another (Allen et al., 1999; Lavasanifar et al., 2002). However, these studies were the first, to our knowledge, that quantitatively compared drug solubilization with !sp values and core-forming block length. We demonstrated that a strong correlation existed between the amount of drug solubilized and the compatibility between that drug and the core-forming block. The PCL block length was also shown to be directly correlated to the amount of hydrophobic drug solubilized. It was suggested that this trend was due to an increase in core volume, providing a larger cargo space for the solubilization of drugs. This relationship was confirmed by the findings in Chapter 4 in which the increased loading capacity of PTX by the nanospheres compared to micelles was ascribed to a larger core volume of the nanospheres possessing a much larger PCL block length. Both nanospheres and micelles displayed controlled in vitro drug release, with release from nanospheres being considerably slower than that from micelles. The decreased release  193 rate of PTX from the nanospheres was likely due to a decreased rate of diffusion caused by an increased interaction of the drug with the longer PCL blocks as well as the higher core microviscosity.  The relationships generated by these studies provide us with the ability to estimate the amount of drug that may be solubilized by one of the copolymers in the series; however, limitations of the studies are apparent. Due to the small number of drugs screened, it is likely that the relationships would not accurately predict the solubilization of any given drug. In order to do so, a large number of compounds would need to be evaluated in order to develop a more appropriate model. For example, particular structural features or properties of drugs may contribute to, or hinder, drug solubilization. Furthermore, these studies only investigated one core-forming block, PCL. Characterization of drug solubilization by a variety of copolymers and the development of models taking into account both the drug and copolymer would ultimately provide a tool which would aid in the appropriate selection of copolymers for drug solubilization.  The physicochemical characterization data gathered in Chapters 2 and 3 were used to select copolymers for further development as PTX nanoparticulate drug delivery systems. It became apparent that the short block copolymers may have limited utility as carrier systems in the blood, due to the higher CMC values and potential for micellar disassembly upon dilution, as well as limited core volume for drug solubilization. Furthermore, the temperature dependent phase behavior may hinder their use, as they are only capable of forming stable colloidal solutions above 30°C. The future of these  194 materials may not be as drug solubilizers, but rather as excipients for the modulation of drug efflux (Zastre et al., 2002; Zastre et al., 2004; Zastre et al., 2007). Copolymers with intermediate PCL block lengths showed more promise as micellar drug delivery systems as they possessed improved thermodynamic stability due to lower CMC values, and increased drug solubilization due to their larger cores. Therefore, of the water soluble copolymers, MePEG114-b-PCL19 was chosen for further evaluation. MePEG114-b-PCL104 was water insoluble, and therefore, did not form micellar dispersions, but rather what we termed nanospheres, characterized by different physicochemical properties compared to micelles. These are discussed in greater depth below.  In Chapter 4, the drug delivery performance characteristics of micelles and nanospheres were compared. Copolymers with short hydrophobic block lengths are readily water soluble and associate into micelles, which are in dynamic equilibrium with free unimers in solution and characterized by a fluid-like core (Riley et al., 2001). Conversely, copolymers with long hydrophobic blocks form larger nanoparticles that are characterized by a more solid-like core, termed nanospheres (Kwon, 1998). However, we recognize that a categorization of amphiphilic copolymer nanoparticles in this manner is an oversimplification. As is it possible to synthesize a wide array of copolymers, ranging in their hydrophilic and hydrophobic block lengths, there will not be a distinct point at which one type of nanoparticle is formed as opposed to the other, but rather, a gradual transition. This has led to a variety of terms used in the literature including crew cut micelles (Zhang and Eisenberg, 1995) and micelle-like nanoparticles (Venkatraman et al., 2005). However, we suggest that the nanoparticles formed from copolymers at either  195 end of the block composition spectrum may be appropriately categorized as either micelles or nanospheres and should possess distinctly different physicochemical properties, which affect drug solubilization, release and distribution in plasma. Accordingly, in these studies, we have classified nanoparticles formed by MePEG114-b- PCL19 as micelles and those composed of MePEG114-b-PCL104 as nanospheres. This classification is supported by the physicochemical characterization data in Chapter 4. The hydrodynamic diameter of nanospheres was larger than that of the micelles demonstrating that the core volume of the nanosphere was considerably greater than that of the micelles. Furthermore, the diameter of the nanospheres increased with the concentration of copolymer in the organic phase used to prepare the particles. Conversely, this copolymer concentration dependence was not observed for the micellar dispersions but rather, the micelles remained at the same diameter, regardless of the concentration of copolymer used to prepare them. The core of the nanospheres was considerably more viscous than that of the micelles, as determined by fluorescence anisotropy. The differences in the physicochemical properties of these two nanoparticles were then related to their drug delivery performance properties and interaction with blood components as discussed below.  In order for a biomaterial to be considered biocompatible, the material and its degradation products should not induce undesirable host reactions (Piskin, 1995). Although determining the biodegradation profiles of the systems were beyond the scope of this study, the hemocompatibilities were assessed by measuring their clotting time and hemolytic potential. Interference with the coagulation pathways was not detected to any  196 significant  extent for either nanoparticle type. This was likely due to a sufficient surface coating of PEG, preventing binding of clotting factors. A small amount of hemolysis was caused by the MePEG114-b-PCL104 nanospheres after 12 hours; however, this was minimal compared to the MePEG17-b-PCL4 control group. The lack of appreciable amounts of membrane disruption by the MePEG114-b-PCL19 micelles and MePEG114-b- PCL104 nanospheres is suggested to be due to a high kinetic stability of the nanoparticles, preventing the intercalation of unimers into the erythrocyte membrane.  As reported in the studies by Ramaswamy et al., we demonstrated that micellar PTX had a similar plasma partitioning profile as free PTX, with the majority of the drug rapidly partitioning into the LPDP fraction (Ramaswamy et al., 1997). Interestingly, when PTX was encapsulated in nanospheres, the majority of the drug remained associated with the copolymer, which separated into its own visible layer. It was speculated that the larger, more solid-like core of the nanospheres prevented the rapid sequestering of the PTX into the plasma protein rich LPDP. This result suggests an advantage for the nanosphere formulation in the retention of PTX in the presence of plasma; however, the results must be interpreted with caution, due to several limitations of the study. There are several possibilities regarding the fate of the micellar PTX. Since the distribution profile of micellar PTX was similar to free PTX, it is possible that the drug was released from the micelles and was associated with the LPDP. It is also conceivable that the drug was still encapsulated in intact micelles and these micelles were associated with the LPDP. Alternatively, it is possible that the micellar structure was disrupted and both the copolymer and the drug were associated with the LPDP. As these studies were not  197 designed to track the disposition of both the copolymer and the PTX, no conclusions can be drawn about whether the drug remained with the micelles or if the micelles were intact. Similarly, even though the majority of PTX was found associated with the nanospheres, this does not necessarily imply that the nanospheres remained separate from the plasma proteins during the incubation period. It is possible that the nanospheres were initially associated with LPDP or lipoproteins and as a consequence of their density, centrifuged to a separate layer.  In conclusion, it was shown in this work that MePEG-b-PCL micelles and nanospheres were capable of increasing the aqueous solubility of PTX. Both nanoparticles displayed in vitro controlled release of PTX and were found to be hemocompatible. It was demonstrated that nanospheres and micelles had different in vitro plasma distribution profiles, indicating that the majority of PTX remained associated with nanospheres. These results suggest that PTX nanosphere delivery systems may lead to prolonged blood circulation and passive targeting of PTX to tumors, therefore, further investigation of the in vivo pharmacokinetics and biodistribution of these systems is warranted.  5.2 SUGGESTIONS FOR FUTURE WORK From the findings of this thesis, a number of future experiments may be proposed. As mentioned above, the design of the PTX plasma partitioning experiments did not allow for the determination of the fate of the copolymers. A logical follow-up to these studies would be to radiolabel the copolymers with 14 C at the MePEG or PCL end. By this method the distribution of the drug and copolymer may be determined enabling us to  198 ascertain whether PTX is retained within the nanoparticles during in vitro and in vivo experiments. The drug distribution studies reported in this work were conducted in plasma; however, it would be required to also investigate the distribution of nanoparticulate PTX in whole blood as the distribution of PTX may be influenced by the presence of erythrocytes. The hemocompatibility of the nanoparticles was assessed by measuring the hemolysis and blood coagulation times; however, further biocompatibility studies may include the investigation of the complement activation and immunogenicity. As previously mentioned, our group has shown that some of the short block length MePEG-b-PCL copolymers reported in this work are capable of reducing the efflux transport of Pgp drug substrates, such as PTX. However, it is evident that micellar systems composed of these short block length copolymers may not be stable enough to be used as a carrier system for PTX. Recent studies have demonstrated that Pgp associated resistance may be overcome by bypassing Pgp efflux via endocytosis of drug loaded nanoparticulates (Chavanpatil et al., 2006). In the current work, it was shown that the MePEG-b-PCL nanospheres displayed enhanced retention of PTX in a stable, solid-like core and thus, may provide an effective vehicle for the delivery of PTX to drug resistant tumours and subsequent endocytotic uptake. Once in the cells, the nanospheres may act as a depot, releasing the encapsulated drug in a controlled manner. Furthermore, Pgp inhibitors, such as the short block length MePEG-b-PCL copolymers may be included in the formulation to prevent PTX efflux once it is released from the nanospheres. These in vitro evaluations would provide the basis for the rational selection of formulations for in vivo investigation of the biodistribution and pharmacokinetics of drug loaded nanoparticles. Ideally such a formulation would achieve prolonged circulation and  199 passive targeting of tumours and would be evaluated for efficacy in a mutidrug resistant cancer animal model.  5.3 REFERENCES  Aliabadi H M, Mahmud A, Sharifabadi A D and Lavasanifar A. Micelles of methoxy poly(ethylene oxide)-b-poly(caprolactone) as vehicles for the solubilization and controlled delivery of cyclosporine A. Journal of Controlled Release 104 (2005) 301-311. Allen C, Maysinger D and Eisenberg A. Nano-engineering block copolymer aggregates for drug delivery. Colloids and Surfaces B: Biointerfaces 16 (1999) 3-27. Burt H M, Zhang X, Toleikis P, Embree L and Hunter W L. Development of copolymers of poly(DL-lactide) and methoxypolyethylene glycol as micellar carriers of paclitaxel. Colloids and Surfaces B: Biointerfaces 16 (1999) 161-171. Chavanpatil M D, Patil Y and Panyam J. Susceptibility of nanoparticle-encapsulated paclitaxel to P-glycoprotein-mediated drug efflux. 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Poly(ethylene oxide)-block-poly(L-amino acid) micelles for drug delivery. Advanced Drug Delivery Reviews 54 (2002) 169-190. Le Garrec D, Gori S, Luo L, Lessard D, Smith D C, Yessine M A, Ranger M and Leroux J C. Poly(N-vinylpyrrolidone)-block-poly(DL-lactide) as a new polymeric solubilizer for hydrophobic anticancer drugs: in vitro and in vivo evaluation. Journal of Controlled Release 99 (2004) 83-101. Nair L S and Laurencin C T. Biodegradable polymers as biomaterials. Progress in Polymer Science 32 (2007) 762-798. Piskin E. Biodegradable polymers as biomaterials. Journal of Biomaterials Science, Polymer Edition 6 (1995) 775-795. Ramaswamy M, Zhang X, Burt H M and Wasan K M. Human plasma distribution of free paclitaxel and paclitaxel associated with diblock copolymers. Journal of Pharmaceutical Sciences 86 (1997) 460-464. Riley T, Stolnik S, Heald C R, Xiong C D, Garnett M C, Illum L, Davis S S, Purkiss S C, Barlow R J and Gellert P R. Physicochemical Evaluation of Nanoparticles Assembled from Poly(lactic acid)-Poly(ethylene glycol) (PLA-PEG) Block Copolymers as Drug Delivery Vehicles. Langmuir 17 (2001) 3168-3174. Venkatraman S S, Jie P, Min F, Freddy B Y C and Leong-Huat G. Micelle-like nanoparticles of PLA-PEG-PLA triblock copolymer as chemotherapeutic carrier. International Journal of Pharmaceutics 298 (2005) 219-232.  201 Zastre J, Jackson J, Bajwa M, Liggins R, Iqbal F and Burt H. Enhanced cellular accumulation of a P-glycoprotein substrate, rhodamine-123, by caco-2 cells using low molecular weight methoxypolyethylene glycol-block-polycaprolactone diblock copolymers. European Journal of Pharmaceutics and Biopharmaceutics 54 (2002) 299- 309. Zastre J, Jackson J and Burt H. Evidence for Modulation of P-glycoprotein-Mediated Efflux by Methoxypolyethylene Glycol-block-Polycaprolactone Amphiphilic Diblock Copolymers. Pharmaceutical Research 21 (2004) 1489-1497. Zastre J, Jackson J K, Wong W and Burt H M. Methoxypolyethylene glycol-block- polycaprolactone diblock copolymers reduce P-glycoprotein efflux in the absence of a membrane fluidization effect while stimulating P-glycoprotein ATPase activity. Journal of Pharmaceutical Sciences 96 (2007) 864-875. Zhang L and Eisenberg A. Multiple morphologies of "crew-cut" aggregates of polystyrene-b-poly(acrylic acid) block copolymers. Science 268 (1995) 1728-1731. Zhang X, Jackson J K and Burt H M. Development of amphiphilic diblock copolymers as micellar carriers of taxol. International Journal of Pharmaceutics 132 (1996) 195-206.  202  APPENDIX 1 ETHICS CERTIFICATE 

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